U.S. patent application number 16/555669 was filed with the patent office on 2020-03-05 for pharmaceutical formulation and system and method for delivery.
The applicant listed for this patent is Rilento Pharma, LLC. Invention is credited to William Andrew Daunch, Raymond A. Dionne, Mark Franklin Hanna, Kevin Neshat, Anthony A. Parker.
Application Number | 20200069595 16/555669 |
Document ID | / |
Family ID | 67953865 |
Filed Date | 2020-03-05 |
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United States Patent
Application |
20200069595 |
Kind Code |
A1 |
Neshat; Kevin ; et
al. |
March 5, 2020 |
PHARMACEUTICAL FORMULATION AND SYSTEM AND METHOD FOR DELIVERY
Abstract
A sustained release pharmaceutical formulation for pain
management comprises an active ingredient, and a water-miscible and
hygroscopic network-forming material, the active ingredient being
dispersed within the water-miscible and hygroscopic network-forming
material. The pharmaceutical may comprise a hydrophobic component,
wherein the active ingredient dispersed within the water-miscible
and hygroscopic network-forming material are together dispersed in
hydrophobic component. Optionally, the pharmaceutical formulation
may be combined with a reinforcing member for providing a system
for sustained release of the pharmaceutical formulation for pain
management.
Inventors: |
Neshat; Kevin; (Raleigh,
NC) ; Daunch; William Andrew; (Cary, NC) ;
Parker; Anthony A.; (Newtown, PA) ; Hanna; Mark
Franklin; (Raleigh, NC) ; Dionne; Raymond A.;
(New Bern, NC) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Rilento Pharma, LLC |
Raleigh |
NC |
US |
|
|
Family ID: |
67953865 |
Appl. No.: |
16/555669 |
Filed: |
August 29, 2019 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
62893413 |
Aug 29, 2019 |
|
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62725694 |
Aug 31, 2018 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61K 9/5089 20130101;
A61L 24/0036 20130101; A61L 2400/04 20130101; A61K 9/0063 20130101;
A61L 24/104 20130101; A61L 2300/402 20130101; A61K 9/5031 20130101;
A61L 2300/622 20130101; A61K 9/5052 20130101; A61L 24/0015
20130101; A61L 2430/12 20130101; A61K 31/445 20130101; A61L 24/046
20130101; A61P 23/02 20180101; A61K 9/7007 20130101; A61L 24/046
20130101; A61L 24/102 20130101; A61L 2300/802 20130101; C08L 67/04
20130101 |
International
Class: |
A61K 9/50 20060101
A61K009/50; A61K 31/445 20060101 A61K031/445 |
Claims
1. A system for sustained release of a pharmaceutical formulation
for pain management, the sustained release system comprising: a
pharmaceutical formulation, including an active ingredient, and a
water-miscible and hygroscopic network-forming material, the active
ingredient dispersed within the water-miscible and hygroscopic
network-forming material; and a reinforcing member.
2. The sustained release system as recited in claim 1, wherein the
active ingredient comprises an anesthetic.
3. The sustained release system as recited in claim 2, wherein the
anesthetic comprises bupivacaine.
4. The sustained release system as recited in claim 1, further
comprising an encapsulating material encapsulating the active
ingredient.
5. The sustained release system as recited in claim 4, wherein the
encapsulating material comprises PLGA.
6. The sustained release system as recited in claim 5, wherein the
encapsulated particles are prepared using a spinning disk or spray
dry atomization process, or an emulsion, solvent extraction
process.
7. The sustained release system as recited in claim 1, wherein the
network-forming material comprises collagen.
8. The sustained release system as recited in claim 1, wherein the
network-forming material comprises gelatin.
9. The sustained release system as recited in claim 8, wherein the
gelatin has a Bloom value of 50 to 325.
10. The sustained release system as recited in claim 1, wherein the
reinforcing member comprises a knitted, woven or non-woven textile,
wherein the interstitial spaces between fibers of the textile are
impregnated with the pharmaceutical formulation.
11. The sustained release system as recited in claim 10, wherein
the textile has a bulk fiber mass per topical unit area of 0.005
g/cm.sup.2 to 0.05 g/cm.sup.2.
12. The sustained release system as recited in claim 1, wherein the
reinforcing member comprises a cellulose hemostatic material.
13. The sustained release system as recited in claim 1, further
comprising a pH modulator.
14. A sustained release pharmaceutical formulation for pain
management, the pharmaceutical formulation comprising: an active
ingredient; a water-miscible and hygroscopic network-forming
material, the active ingredient dispersed within the water-miscible
and hygroscopic network-forming material; and a hydrophobic
component, wherein the active ingredient dispersed within the
water-miscible and hygroscopic network-forming material are
together dispersed in hydrophobic component.
15. The pharmaceutical formulation as recited in claim 14, wherein
the active ingredient comprises an anesthetic.
16. The pharmaceutical formulation as recited in claim 15, wherein
the anesthetic comprises bupivacaine.
17. The pharmaceutical formulation as recited in claim 14, further
comprising an encapsulating material encapsulating the active
ingredient.
18. The pharmaceutical formulation as recited in claim 17, wherein
the encapsulating material comprises PLGA.
19. The pharmaceutical formulation as recited in claim 18, wherein
the encapsulated particles are prepared using a spinning disk or
spray dry atomization process, or an emulsion, solvent extraction
process.
20. The pharmaceutical formulation as recited in claim 14, wherein
the network-forming material comprises collagen.
21. The pharmaceutical formulation as recited in claim 14, wherein
the network-forming polymer comprises gelatin.
22. The pharmaceutical formulation as recited in claim 21, wherein
the gelatin has a Bloom value of 50 to 325.
23. The pharmaceutical formulation as recited in claim 14, further
comprising a reinforcing member.
24. The pharmaceutical formulation as recited in claim 23, wherein
the reinforcing member comprises a knitted, woven or non-woven
textile.
25. The pharmaceutical formulation as recited in claim 23, wherein
the reinforcing member comprises a cellulose hemostatic
material.
26. The pharmaceutical formulation as recited in claim 14, further
comprising a pH modulator.
27. The pharmaceutical formulation as recited in claim 14, wherein
the hydrophobic component is an oil, a wax, or mixtures
thereof.
28. The pharmaceutical formulation as recited in claim 14, wherein
the hydrophobic component is selected from mineral oil, isopropyl
palmitate, caprylic triglyceride, coconut oil, carnauba wax,
beeswax, paraffin wax or mixtures thereof.
29. A pharmaceutical formulation, comprising: 5% to 60% by weight
of an active ingredient; an encapsulated active ingredient, the sum
total of encapsulating material and the encapsulated active
ingredient being 10% to 65% by weight of the pharmaceutical
formulation; 5% to 25% by weight of a water-miscible and
hygroscopic network-forming material; and 15% to 35% by weight of a
hydrophobic component.
30. A system for sustained release of a pharmaceutical formulation
for pain management, the sustained release system comprising: a
pharmaceutical formulation, including 5% to 65% by weight of an
active ingredient, 5% to 25% by weight of a water-miscible and
hygroscopic network-forming material, 20% to 50% by weight of a
hydrophilic component; and up to 15% by weight of a reinforcing
member.
31. A method of delivering a sustained release pharmaceutical
formulation for pain management at a target site of a patient, the
delivery method comprising the steps of: providing a pharmaceutical
formulation, including an active ingredient, a water-miscible and
hygroscopic network-forming material, the active ingredient
dispersed in the water-miscible and hygroscopic network-forming
polymer, and a hydrophobic liquid mixed with the water-miscible and
hygroscopic network-forming polymer including the dispersed
encapsulated active ingredient; and deploying the pharmaceutical
formulation at the target site.
32. A method of delivering a sustained release pharmaceutical
formulation for pain management at a target site of a patient, the
delivery method comprising the steps of: providing a pharmaceutical
formulation, including an active ingredient, and a water-miscible
and hygroscopic network-forming material, the active ingredient
dispersed in the water-miscible and hygroscopic network-forming
polymer, an active ingredient encapsulated in a polymer; blending
water with the water-miscible and hygroscopic network-forming
polymer including the dispersed encapsulated active ingredient; and
deploying the blend at the target site.
Description
CROSS-REFERENCES
[0001] This application is related to U.S. provisional application
No. 62/725,694, filed Aug. 31, 2018, and U.S. provisional
application No. 62/893,413, filed Aug. 29, 2019. The contents of
the provisional applications are incorporated herein by reference
in their entirety, and the benefit of the filing dates of the
provisional applications are hereby claimed for all purposes that
are legally served by such claim for the benefit of the filing
dates.
BACKGROUND
[0002] A pharmaceutical formulation is described and, more
particularly, a sustained release pharmaceutical formulation and a
system and method for delivery of the pharmaceutical formulation
for use, for example, for pain management in wounds such as dental
extractions.
[0003] There is currently no sustained delivery system commercially
available for the specific indication of post-surgical pain after
dental extractions. Ideally, such a product would require minimal
preparation and preferably no preparation by the clinician, it
would be easily placed into the tooth extraction socket or wound
cavity by a clinician, it would have rheological properties that
allow the formulation to be molded to fill the extraction socket or
wound void, it would preferably remain adhered and resist erosion
throughout the treatment duration, it would have no adverse
interactions with blood and would preferably function as a
hemostat, it would have no local (acute or long-term) tissue or
nerve toxicity, it would preferably be comprised of biocompatible
ingredients, it would deliver pain medication both acutely after
surgery and during healing while preferably addressing acute and
sub-acute pain without delaying or adversely affecting wound
healing, and it would preferably enhance wound healing.
[0004] Products that are current benchmarks for rheological
performance in dental surgery and tooth extraction applications
include SURGIFOAM.RTM. Absorbable Gelatin Sponge and SURGIFOAM.RTM.
Absorbable Gelatin Powder, each being examples of sterile porcine
gelatin absorbable sponges or powders intended for hemostatic use
by applying to a bleeding surface ("Surgifoam"). GELFOAM.RTM.
Dental Sponges (absorbable gelatin sponge, USP) is a medical device
also intended for application to bleeding surfaces as a hemostatic.
It is a water-insoluble, off-white, nonelastic, porous, pliable
product prepared from purified pork skin gelatin USP granules and
water for injection, and is able to absorb and hold within its
interstices many times its weight of blood and other fluids.
Gelfoam.RTM. absorbable gelatin powder (absorbable gelatin powder
from absorbable gelatin sponge, USP) is a fine, dry,
heat-sterilized light powder prepared by milling absorbable gelatin
sponge ("Gelfoam"). Soluble collagen powders are another option.
However, compared to Surgifoam and Gelfoam, soluble collagen powder
exhibits a slower rate of gelation since its rate of network
entanglement leads to slower achievement of solidification and
final equilibrium properties. Surgifoam and Gelfoam also have a
significantly higher rate of water adsorption while simultaneously
retaining their solid character; a high overall capacity for water
adsorption; and higher overall compliance with negligible
elasticity at equal water levels in their final equilibrium state.
Commercial collagens generally lead to lower-compliance, rubbery
networks.
[0005] Presently, the pharmaceutical industry is focusing on the
development of sustained release formulations designed to release a
drug at a predetermined rate and to maintain a constant drug level
for a specific period of time with minimal side effects. The basic
rationale behind a sustained release drug delivery system is to
optimize the biopharmaceutical, pharmacokinetic and
pharmacodynamics properties of a drug in such a way that the
utility of the drug is maximized, its side-effects are reduced, and
the disease management goals are achieved. There are several
advantages of sustained release drug delivery over conventional
dosage forms including improved patient compliance due to less
frequent drug administration, reduction of fluctuation in
steady-state drug levels, maximum utilization of the drug,
increased safety margins of potent drugs, and reduction in
healthcare costs through improved therapy and shorter treatment
periods. One of the basic goals of sustained release is to provide
a promising way to decrease the side effects of a drug, first by
preventing the fluctuation of the therapeutic concentration of the
drug in the body, and secondly by reducing the frequency of dose
administration to increase the probability of patient
compliance.
[0006] According to the Centers for Disease Control and Prevention,
drug overdose deaths, including those involving opioids, continue
to increase in the United States. Deaths from drug overdose are up
among both men and women, among all races, and among adults of
nearly all ages. Two out of three drug overdose deaths involve an
opioid. Opioids are substances that work in the nervous system of
the body or in specific receptors in the brain to reduce the
intensity of pain. Overdose deaths from opioids, including
prescription opioids, heroin, and synthetic opioids like fentanyl
have increased almost six times since 1999. In 2017, drug overdoses
of all types averaged 21.7 per 100,000 with opioids alone killing
more than 47,000 people, and with opioids representing 67.8% of all
drug overdose deaths. According to the NIH HEAL Initiative (Helping
to End Addiction Long-term ), more than 25 million Americans suffer
from daily chronic pain. New treatment options for pain are needed
to reduce the number of people exposed to the risks of opioids.
Through the HEAL Initiative, NIH is supporting research to
understand how chronic pain develops, making patients susceptible
to risks associated with opioid use. HEAL is developing a data
sharing collaborative, new biomarkers for pain, and a clinical
trials network for testing new pain therapies. Research efforts are
also focusing on treatments for opioid misuse and addiction.
[0007] According to the American Dental Association's official
policies and statements on substance use disorders including the
opioid crisis, specifically the Statement on the Use of Opioids in
the Treatment of Dental Pain, dentists should follow and
continually review Centers for Disease Control and state licensing
board recommendations for safe opioid prescribing, dentists should
consider treatment options that utilize best practices to prevent
exacerbation of or relapse of opioid misuse, Dentists should
consider nonsteroidal anti-inflammatory analgesics as the
first-line therapy for acute pain management, and dentists should
recognize multimodal pain strategies for management for acute
postoperative pain as a means for sparing the need for opioid
analgesics.
[0008] U.S. Pat. Nos. 8,253,569 and 9,943,466 and U.S. Patent
Application Pub. No. 2018/0169080 describe sustained release
formulations for dental applications. The contents of U.S. Pat.
Nos. 8,253,569 and 9,943,466 and U.S. Patent Application Pub. No.
2018/0169080 are incorporated herein by reference in their
entirety.
[0009] For the foregoing reasons, there is a need for a sustained
release pharmaceutical formulation having rheological behavior
similar to Surgifoam or Gelfoam, and comprising a matrix for
simultaneously achieving and sustaining hemostasis and delivering
active ingredients, such as analgesic or anesthetic drugs to manage
the acute and sub-acute pain during the transition from the
hemostasis phase to the inflammatory phase of wound healing. The
pharmaceutical formulation can be combined with resorbable powders,
fibers or textiles to reinforce the matrix thereby providing a
system for delivering the formulation and for modifying the
rheology so that the formulation adheres to the wound and stays in
place during drug delivery. A reinforcing textile can be foldable
and compressible and have scaffolding and bactericidal properties
as well. Uses of the pharmaceutical formulation and the delivery
system would provide for controlled release of local anesthetic and
anti-inflammatory agents, for example, in a tooth extraction socket
for sustained pain relief from multiple sources of pain and should
promote wound healing. The pharmaceutical formulation should also
satisfy a need to simultaneously address any limits on the
restricted volumes of treatment areas like tooth extraction sockets
while insuring that the formulation has enough mechanical integrity
and cohesive strength to mitigate erosion or detachment from the
wound so that the formulation can deliver the required drug dosage
over time. Ideally, the functional performance and efficacy of the
pharmaceutical formulation and the delivery system with a variety
of drugs should be extendable from the oral surgery model to wounds
or other forms of tissue injury and post-surgical pain.
SUMMARY
[0010] A sustained release pharmaceutical formulation for pain
management is provided. The pharmaceutical formulation comprises an
active ingredient, and a water-miscible and hygroscopic
network-forming material, the active ingredient being dispersed
within the water-miscible and hygroscopic network-forming material.
The pharmaceutical may comprise a hydrophobic component, wherein
the active ingredient dispersed within the water-miscible and
hygroscopic network-forming material are together dispersed in
hydrophobic component. Optionally, the pharmaceutical formulation
may be combined with a reinforcing member for providing a system
for sustained release of the pharmaceutical formulation for pain
management.
[0011] In one aspect, the active ingredient has a weight percent of
less than 60% of the pharmaceutical formulation. The active
ingredient may be present in an acidic form or a basic form. The
active ingredient may comprise an anesthetic. The anesthetic may be
bupivacaine, including an acidic form, a basic form, or a mixture
of acidic and basic forms. Alternatively, the active ingredient is
selected from an analgesic like acetaminophen. Alternatively, the
active ingredient is selected from non-steroidal anti-inflammatory
drugs (NSAID) analgesics. The NSAID may be ibuprofen, naproxen,
meloxicam, ketoprofen, or mixtures thereof. Alternatively, the
active ingredient is a mixture of anesthetics and analgesics.
[0012] The sustained release pharmaceutical formulation and system
may further comprise an encapsulating material encapsulating the
active ingredient. In one embodiment, the encapsulating material is
a polymer, such as PLGA. The PLGA encapsulating material may have
an average particle size of 1 micron to 80 microns, an inherent
viscosity of 0.16 to 1.7 dL/g, a Tg of greater than 37 degrees
Celsius, or a ratio of lactic acid to glycolic acid of 50/50 w/w to
85/15 w/w. The encapsulating material may also comprise an
oligomeric material. The encapsulated particles can be prepared
using a spinning disc spray dry process or an emulsion process.
[0013] In one aspect, the network-forming material has a weight
percent of 5% to 25% of the pharmaceutical formulation. The
network-forming material may comprise a polymer, including either
collagen or gelatin. The gelatin may have a Bloom value of 50 to
325, a viscosity of 1.5 to 7.5 mPa-s, and a mesh value of between 8
and 400.
[0014] In one embodiment, the reinforcing member has a weight
percent of up to 15% of the system. The reinforcing member may
comprise knitted, woven or non-woven fibers, wherein the
interstitial spaces between the fibers are impregnated with the
pharmaceutical formulation. In one aspect, the reinforcing member
comprises a textile, wherein the textile has a bulk fiber mass per
topical unit area of 0.005 g/cm.sup.2 to 0.05 g/cm.sup.2. In
another aspect, the reinforcing member may comprise a cellulose
hemostat material.
[0015] The sustained release pharmaceutical formulation and system
may further comprise a pH modulator. The pH modulator can be an
acid, such as citric acid. The acid has a weight percent of up to
5% of the pharmaceutical formulation. The pH modulator may also be
a base, such as di-sodium citrate. The base has a weight percent of
up to 5%.
[0016] The sustained release pharmaceutical formulation and system
may further comprise a surfactant, an antiemetic, anti-infective,
or chemotherapeutic agent.
[0017] In one aspect, the hydrophobic component is an oil, a wax,
or mixtures thereof. In particular, the hydrophobic component is
selected from mineral oil, isopropyl palmitate, caprylic
triglyceride, coconut oil, carnauba wax, beeswax, paraffin wax or
mixtures thereof.
[0018] In yet another aspect, the water-miscible and hygroscopic
network-forming material does not gel for at least a time period of
24 hours after being suspended within the hydrophobic
component.
[0019] Another embodiment of a sustained release pharmaceutical
formulation for pain management comprises 5% to 60% by weight of an
active ingredient, 10% to 65% by weight of an encapsulating
material in combination with an active ingredient, the
encapsulating material encapsulating the active ingredient, 5% to
25% by weight of a water-miscible and hygroscopic network-forming
material, and 15% to 35% by weight of a hydrophobic component.
[0020] Another embodiment of a system for sustained release of a
pharmaceutical formulation for pain management comprises a
pharmaceutical formulation, including 5% to 60% by weight of an
active ingredient, 10% to 65% by weight of an encapsulating
material in combination with an active ingredient, the
encapsulating material encapsulating the active ingredient, 5% to
25% by weight of a water-miscible and hygroscopic network-forming
material, 20% to 60% by weight of a hydrophilic component, and up
to 15% by weight of a reinforcing member. The hydrophilic component
may comprise glycerin, water, or a mixture thereof.
[0021] A method is also provided for delivering a sustained release
pharmaceutical formulation for pain management at a target site of
a patient. The delivery method comprises the steps of providing a
pharmaceutical formulation, including an active ingredient, a
water-miscible and hygroscopic network-forming material, the active
ingredient dispersed in the water-miscible and hygroscopic
network-forming polymer, and a hydrophobic liquid mixed with the
water-miscible and hygroscopic network-forming polymer including
the dispersed encapsulated active ingredient. The pharmaceutical
formulation is deployed at the target site. The target site may be
a tooth extraction socket.
[0022] Another embodiment of a method of delivering a sustained
release pharmaceutical formulation for pain management at a target
site of a patient comprises the steps of providing a pharmaceutical
formulation, including an active ingredient, and a water-miscible
and hygroscopic network-forming material, the active ingredient
dispersed in the water-miscible and hygroscopic network-forming
polymer, an active ingredient encapsulated in a polymer, blending
water with the water-miscible and hygroscopic network-forming
polymer including the dispersed encapsulated active ingredient, and
deploying the blend at the target site, such as a tooth extraction
socket.
BRIEF DESCRIPTION OF THE DRAWINGS
[0023] For a more complete understanding of the present
formulation, reference should now be had to the embodiments shown
in the accompanying drawings and described below. In the
drawings:
[0024] FIG. 1 is a photograph showing formulation mixtures using
beeswax with three different types of oils (14C-2 with mineral oil,
12019-23-1 with isopropyl palmitate, and 12019-23-2 with caprylic
triglyceride) blended together with powdered bovine gelatin and
PLGA particles, and with each impregnating a textile.
[0025] FIG. 2 is a photograph showing three comparative delivery
systems from FIG. 1 after placing them into the bottom sections of
separate 11 ml glass vials with 2.5 g of added water (representing
the t=0 onset of the pH-neutral water soak experiment at
approximately 20 degrees C.). Formulations from left to right:
12019-23-2, 12019-23-1, and 14C-2.
[0026] FIG. 3 is a photograph showing three comparative delivery
systems from FIG. 1 after placing them into the bottom sections of
separate 11 ml glass vials with 2.5 g of added water (representing
t=24 hours after the onset of the pH-neutral water soak experiment
at approximately 20 degrees C.). Formulations from left to right:
12019-23-2, 12019-23-1, and 14C-2.
[0027] FIG. 4 is a photograph showing three comparative delivery
systems from FIG. 1 after placing them into the bottom sections of
separate 11 ml glass vials with 2.5 g of added water (representing
t=48 hours after the onset of the pH-neutral water soak experiment
at approximately 20 degrees C.). Formulations from left to right:
12019-23-2, 12019-23-1, and 14C-2.
[0028] FIG. 5 is a photograph showing three comparative delivery
systems from FIG. 1 after placing them into the bottom sections of
separate 11 ml glass vials with 2.5 g of added water (representing
t=72 hours after the onset of the pH-neutral water soak experiment
at approximately 20 degrees C.). Formulations from left to right:
12019-23-2, 12019-23-1, and 14C-2.
[0029] FIG. 6 is a photograph showing three comparative delivery
systems from FIG. 1 after placing them into the bottom sections of
separate 11 ml glass vials with 2.5 g of added water (representing
t=120 hours after the onset of the pH-neutral water soak experiment
at approximately 20 degrees C.). Formulations from left to right:
12019-23-2, 12019-23-1, and 14C-2.
[0030] FIG. 7a is a photograph showing hydrophilic system samples
918-1B (left) and 918-1i (right) at t=0 hours after incubation at
37 degrees C. during the pH-2 soak experiment.
[0031] FIG. 7b is a photograph showing hydrophilic system samples
918-1B (left) and 918-1i (right) at t=1.5 hours after incubation at
37 degrees C. during the pH-2 soak experiment.
[0032] FIG. 7c is a photograph showing hydrophilic system samples
918-1B (left) and 918-1i (right) at t=4 hours after incubation at
37 degrees C. during the pH-2 soak experiment.
[0033] FIG. 7d is a photograph showing hydrophilic system samples
918-1B (right) and 918-1i (left) at t=24 hours after incubation at
37 degrees C. during the pH-2 soak experiment.
[0034] FIG. 7e is a scanning electron micrograph of BUP containing
PLGA microspheres produced using the spray drying atomization
method.
[0035] FIG. 7f is an optical microscope image (200.times.
magnification) of BUP containing PLGA microspheres produced using
the emulsion, solvent extraction method.
[0036] FIG. 8 depicts individual UV absorption spectra of fully
dissolved (e.g., GLBG, BUP) and fully dispersed ingredients (e.g.,
PLGA Placebo, BUP encapsulated by PLGA) in pH 2 water at
concentrations that were equivalent to the effective concentrations
used in the fully formulated delivery systems.
[0037] FIG. 9 depicts UV spectra of aliquots removed from the
supernatants of delivery systems comprising hydrophilic components
while soaking in pH-2 water at 37 degrees C.
[0038] FIG. 10a is a photograph showing the hydrophobic
textile-impregnated formulations 14C-3A Placebo, 14C-3B2, and
14C-3A (from left to right) at time=0 hours during the pH-2 soak
experiment at 37 degrees C.
[0039] FIG. 10b is a photograph showing the hydrophobic
textile-impregnated formulations 14C-3A Placebo, 14C-3B2, and
14C-3A (from left to right) at t=1.5 hours during the pH-2 soak
experiment at 37 degrees C.
[0040] FIG. 10c is a photograph showing the hydrophobic
textile-impregnated formulations 14C-3A Placebo, 14C-3B2, and
14C-3A (from left to right) at t=4.0 hours during the pH-2 soak
experiment at 37 degrees C.
[0041] FIG. 10d is a photograph showing the hydrophobic
textile-impregnated formulations 14C-3A Placebo, 14C-3A, and
14C-3B2 (from left to right) at t=24 hours during the pH-2 soak
experiment at 37 degrees C.
[0042] FIG. 10e is a photograph showing the hydrophobic
textile-impregnated formulations 14C-3A Placebo, 14C-3A, and
14C-3B2 (from left to right) at t=4 days during the pH-2 soak
experiment at 37 degrees C.
[0043] FIG. 11a depicts the relative absorbance vs. wavelength for
the hydrophobic delivery system supernatants at t=1.5 hours after
the onset of the water soaking experiments in pH-2 water.
[0044] FIG. 11b depicts the relative absorbance vs. wavelength for
the hydrophobic delivery system supernatants at t=4 hours after the
onset of the water soaking experiments in pH-2 water.
[0045] FIG. 11c depicts the relative absorbance vs. wavelength for
the hydrophobic delivery system supernatants at t=24 hours after
the onset of the water soaking experiments in pH-2 water.
[0046] FIG. 11d depicts the relative absorbance vs. wavelength for
the hydrophobic delivery system supernatants at t=96 hours after
the onset of the water soaking experiments in pH-2 water.
[0047] FIG. 12a depicts the relative absorbance vs. wavelength for
the supernatant of a delivery system created with formulation
14C-3A Placebo, illustrating the progression of the absorbance
curves as a function of time at t=1.5 hours, t=4 hours, t=24 hrs.,
and t=96 hrs. after the onset of the water soaking experiments in
pH-2 water.
[0048] FIG. 12b depicts the relative absorbance vs. wavelength for
the supernatant of a delivery system created with formulation
14C-3B2, illustrating the progression of the absorbance curves as a
function of time at t=1.5 hours, t=4 hours, t=24 hrs., and t=96
hrs. after the onset of the water soaking experiments in pH-2
water.
[0049] FIG. 12c depicts the relative absorbance vs. wavelength for
the supernatant of a delivery system created with formulation
14C-3A, illustrating the progression of the absorbance curves as a
function of time at t=1.5 hours, t=4 hours, t=24 hrs., and t=96
hrs. after the onset of the water soaking experiments in pH-2
water.
[0050] FIG. 13 depicts the time evolution of the absorbance
intensity at 262 nm (i.e., the absorbance maximum for BUP-HCl) for
each of the hydrophilic and hydrophobic formulation delivery
systems
[0051] FIG. 14 displays a relative absorbance vs. time comparison
of placebo devices 14C-3E (with citric acid) and 14C-3A (without
citric acid).
[0052] FIG. 15 illustrates the relative BUP concentration (mg/ml)
vs. time (hrs.) as estimated from the UV absorption spectra of the
supernatants that were sampled during the time evolution of the
pH-2 water-soak experiments.
[0053] FIG. 16 illustrates the relative rates of BUP elution
(mg/ml/hour) together with the data ranges used for establishing
the best linear fitting parameters.
[0054] FIG. 17 illustrates the relative rates of BUP elution with
the [BUP] expressed in terms of the fraction of eluted
BUP=[BUP]/[BUP].sub.theoretical=[BUP]/17.14.
DESCRIPTION
[0055] A sustained release pharmaceutical formulation and system
and method for delivery of the pharmaceutical formulation for, for
example, pain management are described. The pharmaceutical
formulation comprises an active ingredient optionally encapsulated
in an encapsulant, a water-miscible and hygroscopic network-forming
material, and, optionally, a reinforcing member. Embodiments of the
pharmaceutical formulation and system and method include: 1) those
comprising a dry powder mixture, including components that are
first mixed as powders and then hydrated and masticated before end
use; 2) those that are formulated with hydrophobic components and
then hydrated before end use; 3) those that are formulated with
hydrophobic components and then allowed to hydrate in vivo; 4)
those that are formulated with hydrophobic components and then
impregnated into the reinforcing member and hydrated and masticated
before end use; 5) those that are formulated with hydrophobic
components and then impregnated into the reinforcing member and
allowed to hydrate in vivo; and 6) those that are formulated with
either hydrophobic or hydrophobic components and then mixed with
reinforcing members that are powders, fibers or granulated
textiles, then hydrated and masticated before end use or allowed to
hydrate in vivo. The reinforcing member may be reinforcing oxidized
regenerated cellulose (ORC) or carboxymethyl cellulose sodium (CMC)
powder or fibers, or impregnated knitted, woven or non-woven ORC
and CMC textiles. The impregnated textile functions as a delivery
system and provides a cost-effective, manufacturing-effective, and
clinically advantageous set of options for retaining the
formulation within the tooth extraction socket.
[0056] The network-forming material, like gelatin or others, is
required in certain embodiments to act as a binder for the
dispersed ingredients, particularly upon hydration of the
pharmaceutical formulation to deter macroscopic phase separation
and erosion during deployment and hydration. Upon hydration of the
formulation, either in vivo or alternatively ex vivo via
mastication with water prior to use, it is believed that
phase-inversion occurs whereby the network-forming material or
cellulose textile becomes a plasticized and entangled network that
serves as a binder for the encapsulated active-ingredient particles
as well as for other dispersed ingredients. Simultaneously, the
hydrophobic components (e.g., oil, wax), remain dispersed within
the hydrated matrix and resist undergoing macroscopic phase
separation and exudation. The post-hydration binding capacity that
is provided by the plasticized network is necessary to prevent
premature erosion of the formulation from the dental extraction
socket or wound. The state of the dispersion and the degree of
gelatin aggregation throughout these phase-inversion transformation
processes will have an impact on the time-dependent release profile
of active ingredients.
[0057] In an alternative embodiment, the pharmaceutical formulation
may be prepared without the use of the network-forming material,
provided that the textile material is capable of becoming a binder
for the dispersed encapsulated active ingredient when the
formulation is hydrated. Upon hydration, either in vivo or
alternatively ex vivo via mastication with water prior to use, it
is hypothesized that phase-inversion occurs whereby the
network-forming material, the reinforcing member, or both become
plasticized and serve as a binder for the encapsulated active
ingredient particles. The binding is necessary to prevent premature
erosion of the pharmaceutical formulation from the dental
extraction socket or wound. The state of the dispersion and the
degree of aggregation throughout these transformation processes has
an impact on the release profile of the active ingredient. Thus,
the state of dispersion is an important factor that will impact the
release profile. However, the key to consistent release performance
will not necessarily be in achieving an aggregate-free state of
dispersion. Instead, the key to release performance will be in
achieving reproducibility and consistency for any given state of
dispersion that simultaneously satisfies manufacturing constraints
and end use performance targets.
[0058] The various embodiments of the pharmaceutical formulation
have certain morphological and functional attributes in common.
Namely, each embodiment is functionally capable of undergoing in
vivo hydration. Each embodiment facilitates controlled time release
delivery of active ingredient when deployed in fixed-volume
applications, such as within dental extraction sockets. Each
embodiment is capable of inter-mixing with oral fluids such as
saliva and blood in vivo to yield homogeneous structures that
remain cohesively intact for sustained periods of time, thus
enabling each embodiment to perform simultaneously as hemostats and
as sustained release devices. Each comprises a network-forming
material as a binder phase that serves as a matrix for suspending
particulates, including encapsulated microparticles, such as
poly(lactic-co-glycolic acid) (PLGA) encapsulated bupivacaine
(BUP). Moreover, each binder phase may further comprise a liquid
carrier that modulates the rheo-mechanical characteristics of the
pharmaceutical formulation.
[0059] Although the various embodiments of the formulation have
many global similarities, there are also several important
distinctions. One of the most important distinctions stems from the
compositional and physico-chemical differences in the components
that constitute each of their respective binder phases. For liquid
components, the polarity of the compounding liquid and the
propensity for the liquid carrier to cause gelation of gelatin are
the delineating factors for the categorization. The recognition of
the importance of this seemingly minor distinction is one that has
facilitated the creation of several distinct embodiments, each
having different structural and functional features.
[0060] An embodiment of the pharmaceutical formulation is
compounded with a high polarity liquid, wherein the liquid is one
that induces gelation of gelatin prior to the deployment of the
formulation. A compliant dough-like material is formed that can be
deployed for in vivo drug delivery. When the choice of polar liquid
is water or a water solution, the formulation preferably takes the
form of a pre-packaged dry-powder mixture that is hydrated prior to
deployment. When the choice of the high-polarity liquid is one that
is more conducive to shelf-stability, such as glycerin or a high
polarity liquid solution such as glycerin and water, a compliant
dough-like material is formed that can be deployed as a stand-alone
device for in vivo drug delivery. The mixture can be compounded
during manufacturing with the high polarity liquid to form a
compliant dough-like material and packaged as a compliant,
formable, shelf-stable device that can be directly deployed in end
use environments without the need for mixing with water or saline
solution. The preferred high polarity liquids for this application
are biostable and resist microbial growth during storage. Although
these types of formulations can be optionally mixed and hydrated
with water if so desired, they are unique in that they can be
directly deployed for in vivo hydration. These formulations can
also be optionally reinforced with fibrous materials, such as
knitted, woven, or non-woven cellulose textiles including
hemostats, to form a composite like structure.
[0061] An embodiment of the pharmaceutical formulation is
compounded with a low polarity liquid, wherein the liquid is one
that does not induce premature gelation of gelatin prior to the
deployment of the formulation. This embodiment of the
pharmaceutical formulation is compliant, formable, shelf-stable and
can be directly deployed in end use environments without the need
for premixing with water or saline solution. Although these types
of formulations can be optionally premixed and pre-hydrated with
water if so desired, they are unique in that they can be directly
deployed for in vivo hydration. These formulations can also be
optionally reinforced with fibrous materials, such as knitted,
woven and non-woven cellulose fiber textiles including
hemostats.
[0062] Embodiments of the delivery system, wherein a pharmaceutical
formulation is reinforced with a fibrous material to form a
composite like structure, can also be packaged for deployment and
then subsequently deployed for in vivo hydration. The fibrous
component can be either knitted, woven or non-woven, but a
particularly advantageous type of fibrous component for this
purpose is a low knit density cellulose hemostat knitted textile,
which when impregnated with the pharmaceutical formulation
positively enhances the formulation by increasing its strength, its
durability, and its functionality during deployment. These types of
delivery systems can be optionally hydrated with water, but they
are uniquely acceptable for direct deployment and for subsequent in
vivo hydration. The delivery systems tend to resist erosion, and
they can be used to achieve controlled time-release delivery
profiles of active ingredients like bupivacaine over periods of
multiple days.
[0063] In each of the embodiments, the pharmaceutical formulation
is designed to co-disperse network-forming material together with a
variety of other ingredients, including for example, either
unimodal, bi-modal or tri-modal particle size distributions of
active ingredients, particulates of active ingredients encapsulated
by an encapsulating material, or mixtures thereof.
[0064] In one embodiment, the encapsulating material may comprise a
polymer. Polyanhydrides and polyesters are two classes of polymers
often used for controlled release purposes. Polyanhydrides are a
class of polymers composed of hydrolytically labile anhydride
linkages that can be easily modified by vinyl moieties or imides to
create cross-linkable systems, permitting the tailoring of release
rates to the degree of cross-linking density. Mass loss of
polyanhydrides follows a surface degradation mechanism, and drug
release is exclusively controlled by surface erosion processes.
Polyesters such as poly(.epsilon.-caprolactone) (PCL), poly(lactic
acid) (PLA), and poly(lactide-co-glycolide) (PLGA) have been used
in controlled-release formulations currently approved by the FDA.
Among these polymers PLGA is one of the most studied diblock
copolymers for microencapsulation. Unlike polyanhydrides, PLGA
undergoes bulk erosion, with drug release occurring by both
diffusion and erosion processes. The drug release kinetics are
influenced by the several characteristics of the PLGA polymer,
including copolymer composition, molecular weight, crystallinity,
and drug-polymer interactions. In addition to polyanhydrides and
polyesters, microparticles made from copolymers of polyanhydrides
and polyesters have also been investigated for their ability to
achieve better controlled release of drugs.
[0065] The polymer polylactic-co-glycolic acid (PLGA) is an
encapsulant that is well known in the art. With PLGA, the higher
the percentage of lactide units, the longer the polymer lasts
before degrading in the presence of water. In addition, the higher
the molecular weight of PLGA, the greater the mechanical strength.
The degradation rates of PLGA can be influenced by different
parameters including, for example, (i) the molecular weight,
whereby degradation rates have been reported to range from several
weeks to several months with increasing molecular weights ranging
from 10-20 to 100 kDa; (ii) the ratio of glycolic acid (GA) to
lactic acid (LA), whereby PLGA with a higher LA contents are less
hydrophilic, absorb less water and subsequently degrade more slowly
as a consequence of the presence of methyl side groups in poly-LA
making it more hydrophobic than poly-GA (one exception to this rule
being the 50:50 copolymer which exhibits faster degradation); (iii)
stereochemistry, whereby mixtures of D and L lactic acid monomers
are most commonly used for PLGA fabrication because the rate of
water penetration is higher in amorphous D,L regions, leading to
accelerated PLGA degradation; and (iv) end-group functionalization,
whereby polymers that are end-capped with esters, as opposed to the
free carboxylic acid, demonstrate longer degradation half-lives. In
addition, the geometric shape of the reinforcing member will
strongly affect PLGA degradation behavior by influencing the
accessibility of water. It has also been reported that acidic
surrounding media will accelerate PLGA degradation due to
catalysis.
[0066] The glass transition temperature (Tg) of PLGA is reported to
be above 37.degree. C., thereby providing PLGA with polymer chain
rigidity and macro rigidity under ambient conditions and at body
temperature. Further, it has been noted that Tg of PLGA decreases
with decreasing LA content, and with decreasing molecular
weight.
[0067] PLGA copolymers are commercially available with various LA
to GA ratios, including 50/50, 65/35, 75/25, and 85/15; with glass
transition temperatures ranging from 45 to 55 degrees C.; with
inherent viscosities ranging from 0.55 to 0.75 dL/g; with tensile
strengths ranging from 6000 to 8000 psi; with elongations ranging
from 3 to 10%; and with modulus values ranging from
2.times.10.sup.4 to 4.times.10.sup.4 psi. These products are also
described as having degradation/resorption time windows that
generally increase with increasing LA contents. PLGA having LA/GA
ratios of 65/35 degrade in about 3-4 months, LA/GA ratios of 75/25
degrade in about 4-5 months, LA/GA ratios of 85/15 degrade in 5 to
6 months, and where ratios of 50/50 (the exception) degrade in
about 1-2 months. Resomer RG504 available from Evonik (a
poly(D,L-lactide-co-glycolic acid) copolymer with LA/GA=50/50, CAS
#26161-42-2) is reported to have an inherent viscosity (IV) of 0.4
to 0.6 dL/g, a Tg of 46-50 degrees C., a molecular weight of
38,000-54,000 amu, and a degradation timeframe of less than 3
months. Other types of D,L-PLGA copolymers available from Evonik
that are suitable for use in making devices of the types described
herein include those with LA/GA ratios of 50/50 with IV ranging
from 0.16 to 0.74; LA/GA ratios of 65/35 with IV ranging from 0.32
to 0.44; LA/GA ratios of 75/25 with IV ranging from 0.16 to 1.2;
and LA/GA ratios of 85/15 with IV ranging from 1.3 to 1.7.
[0068] For the present sustained release formulation, suitable PLGA
copolymer are amorphous types with LA/GA ratios ranging from 50/50
to 85/15, with IV values ranging from 0.16 to 1.7, and with Tg
values ranging from 37 to 60 degrees C. More preferably, PLGA
copolymers will include those with LA/GA ratios ranging from 50/50
to 75/25, with IV values ranging from 0.16 to 0.75, and with Tg
values ranging from 40 to 55 degrees C.
[0069] In addition, materials other than PLGA polymers may also be
used as encapsulants, such as naturally derived and synthetic
polymers and oligomers. Preferred naturally derived encapsulants
include carbohydrate polymers such as plant derived starch and
starch derivatives, cellulose and cellulose derivatives; plant
exudates such as gum arabic, gum karaya and mesquite gum; plant
extracts such as galactomannans and soluble soybean;
polysaccharides; marine derived carrageenan and alginate;
microbial/animal derived xanthan, gellan, dextran, hyaluronic acid
(natural and cross-linked), albumin, collagen, gelatin and
chitosan; plant proteins such as gluten and isolates from pea and
soy; microbial/animal derived proteins including caseins, whey
proteins and gelatin; and plant and animal derived lipids including
fatty acids, alcohols, glycerides, waxes such as carnauba wax and
beeswax, and phospholipids. Preferred synthetic encapsulants
include homopolymers of polyester-based synthetic polymers like
poly (.epsilon.-caprolactone) (PCL), poly(glycolic acid) (PGA),
poly (lactic acid) (PLA), and poly(phosphoesters) (PPE);
poly(ethylene glycol) (PEG), also known as polyethylene oxide
(PEO), Poly(2-oxazolines) (POX), polyvinyl alcohol (PVA),
poly(N-vinylpyrrolidone) (PVP), blends of polyvinyl acetate (PVAc)
and povidone (PVP), as well as diblock and triblock copolymers and
graft polymers of the aforementioned. Other microencapsulant
material examples can include hydrophobic materials coated via
fluid bed technologies, such as paraffin wax, fractionated palm
oil, hydrogenated palm oil, mono and diglycerides, hydrogenated
cottonseed oil, hydrogenated soybean oil, hydrogenated castor oil,
beeswax, carnauba wax, and distilled monoglycerides; aqueous-based
coatings such as hydroxypropyl methylcellulose (HPMC), gums,
poly(vinyl alcohol) polymers and copolymers, poly(vinyl
pyrrolidone) polymers and copolymers, cellulose polymers,
poly(maleic anhydride) polymers and copolymers, including acid
forms, anhydride forms, acid salt forms, and mixtures thereof,
collagens; and solvent-borne coatings such as ethyl cellulose
dissolved in an alcohol. Other examples of natural and synthetic
polymers known to those skilled in the art can include
carbohydrates such as starch, modified starches, dextrins, sucrose,
cellulose and chitosan; gums such as arabic gum, alginate and
carrageenan; lipids such as wax, paraffin, monoglycerides and
diglycerides, hydrogenated oils and fats; inorganic materials such
as calcium sulfate and silicates; and proteins such as gluten,
casein, gelatin and albumin; each employing encapsulation methods
such as, spray drying, spray cooling, extrusion, coacervation,
lyophilization, and emulsification (da Silva, P. T., et al,
"Microencapsulation: concepts, mechanisms, methods and some
applications in food technology," Ciencia Rural, Santa Maria, v.
44, n. 7, p. 1304-1311, July, 2014).
[0070] PLGA microspheres or microspheres made from the
aforementioned materials can be manufactured by many methods of
microencapsulation, incorporating active ingredients for the
purpose of modulating drug delivery. There are preferred techniques
that emphasize processes that have produced commercially
significant products such as: coacervation; interfacial and in vivo
polymerization; single and double emulsion techniques such as
solvent evaporation, solvent extraction and cross-linking emulsion;
supercritical fluid techniques such as rapid expansion of
supercritical solution (RESS) and supercritical fluid anti-solvent
crystallization (SAS) processes; spray drying; spray coating;
centrifugal extrusion; and rotational suspension separation.
[0071] Active ingredients for pain management may include an
anesthetic or mixture of anesthetics to reduce the sensation of
pain in the area to which they are applied. These anesthetics can
be formulated alone, as mixtures and can be combined with an
anesthetic vehicle like water, a vasoconstrictor like epinephrin, a
reducing agent like sodium metabisulfite, preservatives like methyl
paraben, and buffers. Anesthetics can be amino esters such as
amylocaine, ambucaine, benzocaine, butacaine, chloroprocaine,
cocaine, cyclomethycaine, demethocaine (Larocaine), piperocaine,
propoxycaine, procaine (novocaine), proparacaine and tetracaine
(amethocaine). Anesthetics can also be amino amides such as
articaine, bupivacaine, cinchocaine (dibucaine), etidocaine,
levobupivacaine, lidocaine (lignocaine), mepivacaine, prilocaine,
ropivacaine and trimecaine. Anesthetics can also come from
naturally derived sources. Terpenoids, alkaloids and flavonoids are
anesthetic agents of plant origin because they meet the mechanistic
requirements to interact with receptors, channels and membranes.
Naturally derived anesthetics include saxitoxin, neosaxitoxin,
tetrodotoxin, thymol, menthol, eugenol, cocaine, spilanthol,
capsaicin, eunal, propinal, propandid and propofol. Anesthetics as
active ingredients can be racemic mixtures, or the R or S isomers
of the anesthetic depending on absorption, distribution, potency,
toxicity and therapeutic action requirements. Anesthetics as active
ingredients can be the free base form or the ionized form as a
hydrochloride salt.
[0072] Active ingredients for pain management may include
analgesics like acetaminophen and ziconotide, that provide relief
from pain without causing sleep or loss of consciousness.
[0073] Analgesics can be from the class of salicylates such as
magnesium salicylate, aspirin, choline salicylate/magnesium
salicylate, diflunisal, salsalate, aspirin/citric acid/sodium
bicarbonate.
[0074] Analgesics can be from the class of nonsteroidal
anti-inflammatory drugs (NSAIDS) such as ketoprofen, fenoprofen,
tolmetin, diclofenac/misoprostol, piroxicam, sulindac,
indomethacin, diclofenac, etodolac, ibuprofen, flurbiprofen,
ketorolac, naproxen, meloxicam, diflunisal, esomeprazole/naproxen,
famotidine/ibuprofen, mefenamic acid, oxaprozin, nabumetone,
bromfenac, and meclofenamate.
[0075] Analgesics can be from the class of Calcitonin gene-related
peptide (CGRP) inhibitors such as fremanezumab, erenumab,
galcanezumab and Eptinezumab.
[0076] Analgesics can be from the class of Cyclooxygenase-2 (Cox-2)
inhibitors such as amlodipine, valdecoxib and celecoxib.
[0077] Analgesics can be from the class of antimigraine agents such
as frovatriptan, acetaminophen/dichloralphenazone/isometheptene
mucate, almotriptan, caffeine/ergotamine naproxen/sumatriptan,
rizatriptan, naratriptan, eletriptan, sumatriptan, zolmitriptan,
dihydroergotamine, and ergotamine.
[0078] Analgesics can be from the class of narcotics, such as
meperidine, opium, methadone, hydromorphone, codeine, fentanyl,
oxycodone, oxymorphone, nalbuphine, morphine, butorphanol,
levorphanol, buprenorphine, propoxyphene, tramadol, tapentadol,
pentazocine, hydrocodone, alfentanil, remifentanil, and
sufentanil.
[0079] Although narcotic analgesics may be employed, non-narcotic
types are preferred. If narcotic types are used, it is preferable
that they be of the localized type, capable of agonizing localized
neuroreceptors for localized pain relief, and incapable of crossing
the blood brain barrier so as to minimize possible tendencies for
addiction.
[0080] Analgesics can be combined to contain at least one analgesic
in combination with another medicine or medicines, and when
combined generally have different ways of working to relieve pain,
such as acetaminophen/caffeine/magnesium salicylate,
aspirin/meprobamate acetaminophen/butalbital,
acetaminophen/caffeine, acetaminophen/caffeine/isometheptene
mucate, acetaminophen/pamabrom/pyrilamine, aspirin/diphenhydramine,
acetaminophen/pamabrom, acetaminophen/butalbital/caffeine,
aspirin/butalbital/caffeine, acetaminophen/aspirin,
acetaminophen/phenyltoloxamine,
acetaminophen/aspirin/caffeine/salicylamide, aspirin/caffeine,
acetaminophen/aspirin/caffeine, acetaminophen/caffeine/pyrilamine,
acetaminophen/diphenhydramine, diphenhydramine/naproxen,
diphenhydramine/ibuprofen, aspirin/caffeine/salicylamide,
acetaminophen/magnesium salicylate/pamabrom,
acetaminophen/phenyltoloxamine/salicylamide,
acetaminophen/pyrilamine, and diphenhydramine/magnesium salicylate.
Narcotic and non-narcotic analgesic combinations include
belladonna/opium, aspirin/butalbital/caffeine/codeine,
meperidine/promethazine, acetaminophen/butalbital/caffeine/codeine,
ibuprofen/oxycodone, acetaminophen/pentazocine,
hydrocodone/buprofen, buprenorphine/naloxone,
acetaminophen/oxycodone, acetaminophen/caffeine/dihydrocodeine,
acetaminophen/hydrocodone, naloxone/pentazocine,
acetaminophen/tramadol, acetaminophen/propoxyphene,
aspirin/oxycodone, naloxone/oxycodone, acetaminophen/codeine,
morphine/naltrexone, acetaminophen/benzhydrocodone,
aspirin/caffeine/dihydrocodeine, and naltrexone/oxycodone.
[0081] Active ingredients of these aforementioned types may also be
optionally employed without the use of a polymer microencapsulant,
blending them directly into the network forming matrix. Mixed types
of microencapsulated and non-encapsulated types can also be
employed.
[0082] Other types of active ingredients can also be included as
encapsulated on non-encapsulated adjuncts to satisfy a number of
medical purposes, including for example, anti-infectives,
antiemetics, and chemotherapeutic agents.
[0083] Anti-infectives describe any medicine that is capable of
inhibiting the spread of an infectious organism or by killing the
infectious organism outright, encompassing antibiotics,
antifungals, anthelmintics, antimicrobials, antimalarials,
antiprotozoals, antituberculosis agents, and antivirals. In
addition to the aforementioned active ingredients for pain
management, antibiotic, antimicrobial and antifungal
anti-infectives are preferred adjunct active ingredients.
Antibiotics such as penicillin, amoxicillin, amoxicillin/clavulanic
acid, clindamycin, azithromycin, and metronidazole are preferred
adjunct active ingredients. Antifungals such as fluconazole,
clotrimazole, nystatin, itraconazole, and amphotericin B are
preferred adjunct active ingredients.
[0084] Antiemetics are drugs that are effective against vomiting
and nausea. Antiemetics are typically used to treat the side
effects of opioid analgesics, general anesthetics, and cancer
chemotherapy. In addition to the aforementioned active ingredients
for pain management, antiemetic drugs for post-surgical nausea such
as dexamethasone, droperidol, granisetron, metoclopramide, and
ondansetron are preferred adjunct active ingredients. Antiemetic
drugs for chemotherapy nausea (e.g., chemotherapy for treating head
and neck cancers) such as aprepitant, dexamethasone, dolasetron,
granisetron, ondansetron, palonosetron, prochlorperazine,
rolapitant, and cannabinoids are preferred adjunct active
ingredients.
[0085] Chemotherapeutic agents, also referred to as antineoplastic
agents, are used to directly or indirectly inhibit the
proliferation of rapidly growing cells, typically in the context of
malignancy. They are classified according to their mechanism of
action and include alkylating agents, antimetabolites,
topoisomerase inhibitors, and mitotic inhibitors. In addition to
the aforementioned active ingredients for pain management, for
cancer that arises in the head or neck region (in the nasal cavity,
sinuses, lips, mouth, salivary glands, throat, or larynx),
chemotherapeutic agents such as bleomycin sulfate, cetuximab,
docetaxel, erbitux (Cetuximab), Hydrea (Hydroxyurea), Hydroxyurea,
Keytruda (Pembrolizumab), Methotrexate, Nivolumab, Opdivo
(Nivolumab), Pembrolizumab, Taxotere (Docetaxel), and Trexall
(Methotrexate) are preferred adjunct active ingredients
[0086] Pharmacokinetic modulating additives can be optionally
encapsulated or used directly in the formulation mixture, for
example, citric acid, ascorbic acid, palmitic acid, dodecanedioic
acid, sebacic acid, fatty acids such as stearic acid, oil-soluble
types or water-soluble types, to influence the conversion of
anesthetic free base its respective acid form. Additives can be
optionally encapsulated or used directly in the formulation mixture
to prolong the duration of anesthetic analgesia, for example
epinephrine, clonidine, dexmedetomidine, buprenorphine,
dexamethasone, tramadol, sodium bicarbonate, and midazolam. Many
materials are suitable for use as the water-miscible and
hygroscopic network-forming component in the present pharmaceutical
formulation. Hygroscopic network-forming polymer components can
include soluble collagen and gelatin; tree exudates of which
arabic, ghatti, karaya, and tragacanth are examples; seaweed
colloids including agar, agarose, Irish moss, carrageenin, and
alginates as examples; extracts from seeds of locust bean, locust
kernel, and quince seed gums as examples; manufactured and modified
dextrins; water-dispersible or water-soluble derivatives of
cellulose; and the like. These types of hygroscopic network forming
polymers can also be used as encapsulants for various active
ingredients if so desired. In such cases, the encapsulant serves
two purposes: it encapsulates the active ingredient to form a
diffusion barrier; and it provides the capacity to form an
entangled network when the device is hydrated either prior to end
use, or in vivo.
[0087] Other types of synthetic water-miscible and hygroscopic
network-forming components can also be employed. For example,
poly(maleic anhydride) polymers and copolymers, including acid
forms, anhydride forms, acid salt forms, and mixtures thereof are
particularly useful for producing networks with varying degrees of
water miscibility, varying degrees of erosion resistance, varying
degrees of capacity for adhesion to membrane tissue, and varying
levels of compliance in their hydrated state. One example of a
class of such copolymers includes the free acid and anhydride forms
of poly(maleic anhydride-co-vinyl methyl ether) (PMAVE). In its
free acid form, the polymer has greater water miscibility, and
exhibits higher tissue membrane adhesion characteristics. Water
miscibility, solubility and adhesion can be controlled through a
combination of factors, including for example, by controlling the
mole ratio of free acid to anhydride within the copolymer, and by
controlling the molecular weight and molecular weight distribution
of the copolymer. In addition, by selective use of monovalent and
divalent counterions, salts of the various types of free-acid
copolymers can be formed, including for example, monovalent Na
salts, di-valent Ca salts, di-valent Mg salts, and mixtures
thereof. The rate of water ingress and the degree of water
miscibility with these types of polymers increases with increasing
mole % of free acid, and decreases with increasing acid salt
complexation, and with increasing valency of the counterion, where
Na salts are most soluble, and Ca and Mg salts are less soluble.
The mechanical compliance characteristics of such polymers are also
known to increase with increasing mole percentages of free acid,
and to decrease with increasing mole percentages of cation
complexation, and also with increasing valency of the counterion.
With these types of controlling levers, including the mole ratio of
free acid form to salt form to anhydride form, the mole ratio of Na
to Ca to Mg counterions, the average molecular weights and
molecular weight distributions of the polymer types or mixtures
thereof, it is possible to create a broad range of mechanical
properties, adhesive properties, water miscibility characteristics,
and network forming properties.
[0088] Gelatin is classified as a mixture of water-soluble proteins
of high average molecular weights, also present in collagen. The
proteins are extracted by boiling skin, tendons, ligaments, bones,
etc. in water. Type A gelatin is derived from acid-cured tissue and
Type B gelatin is derived from lime-cured tissue. Below
35-40.degree. C. gelatin swells in and absorbs 5-10 times its
weight of water to form a gel. Gelatin is soluble in glycerol and
acetic acid, and more soluble in hot than in cold water. It is
practically insoluble in most organic solvents such as alcohol,
chloroform, carbon disulfide, carbon tetrachloride, ether, benzene,
acetone, and oils.
[0089] Bloom is a characteristic used to describe gelatin referring
to gel strength. Bloom is related to molecular weight and is
therefore a factor that affects the mechanical elasticity of
gelatin in its plasticized, gelled state. Bloom tests can be
conducted using a standardized measurement (e.g., the force
required to depress a prescribed area of the surface of a 6.67%
gelatin gel at 10.degree. C. (50.degree. F.) to a distance of 4
mm). The bloom values for one family of commercial gelatin brands
from Rousselot.RTM. are reported to range from 75 to 300 grams. As
such, the gelatins are classified as follows: 1) High bloom--gel
strength above 200 grams; 2) Medium bloom--gel strength between 120
and 200 grams; and 3) Low bloom--gel strength less than 120 grams.
There is a general relationship between bloom and average molecular
weight, where Bloom number generally correlates with average
molecular weights as follows: 50-125 (Low Bloom)=20,000-25,000 amu;
175-225 (Medium Bloom)=40,000-50,000 amu; and 225-325 (High
Bloom)=50,000-100,000 amu.
[0090] A number of gelatin types can be employed in the sustained
release pharmaceutical formulation, including porcine, bovine,
piscine, vegetable, type-A, type-B, or mixtures thereof.
Commercially available matrix proteins, for example Surgifoam and
Gelfoam, may also be used. The bloom values may range from 50 grams
up to 325 grams depending on the desired rate of fluid uptake and
the desired mechanical compliance for the device. Gelatins with
higher bloom values are generally slower to adsorb water and will
lead to lower compliance when they are gelled. In addition, gelatin
types having different bloom values can be mixed at different
weight ratios to achieve intermediate water-uptake rates and
intermediate compliance characteristics. Desirable properties of
the sustained release pharmaceutical formulation can be achieved
with bloom values ranging from about 50 to 325, but preferably from
100 to 300, and more preferably from about 150 to 250.
[0091] Viscosity is also an important factor that affects the
rheological behavior of gelatin solutions. Once dissolved in water,
gelatins with bloom values covering the aforementioned range will
yield solutions having viscosities typically ranging from 1.5 to
7.5 mPa-s. Viscosity is measured by a standardized method whereby
the flow time of 100 ml of a 6.67% gelatin solution at 60.degree.
C. (140.degree. F.) is measured when the solution is passed through
a standard pipette. Desirable properties of the pharmaceutical
formulation can be achieved with viscosity values preferably
ranging from about 1.5 to 7.5 mPa-s, and more preferably from about
3 to 6.5 mPa-s.
[0092] Particle size distribution is another important physical
attribute for the sustained release pharmaceutical formulation.
Generally, the larger the particle size (smaller mesh size), the
lower the viscosity of the resulting dispersion at constant weight
ratios of particle to carrier. This factor can be represented by
the mesh size of standard screens that are used for testing
particle size distributions of particulate materials. A single
positive mesh value is interpreted to mean the mesh value at which
90% by weight of the particulates are retained by the mesh screen
when a distribution of particulates is passed through the mesh. For
example, a reported mesh value of 30 (corresponding to a particle
size of about 0.6 mm) would indicate that 90% by weight of the
particle size distribution is retained by a mesh 30 screen when a
distribution is passed through the screen, further indicating that
90% by weight of the distribution contains particulates that are
0.6 mm or larger. For the case of Rousselot.RTM. brands of gelatin,
products are reported to included 8 mesh (2.36 mm) and 18 mesh
(1.00 mm) at the upper range, and 30 mesh (0.60 mm) and 60 mesh
(0.25 mm) at the lower range. The sustained release pharmaceutical
formulation can be adjusted with a variety of 90% mesh particle
sizes ranging from about 400 mesh or higher (0.037 mm or lower) to
about 8 mesh (2.36 mm). Particle size distributions and hence
vehicle rheology and fluid uptake rates can be further adjusted by
blending distributions with different mesh values (e.g., 350 mesh
blended with 60 mesh) and at varying weight ratios to yield
rheological characteristics and fluid uptake rates that are
commensurate with the end use needs for the application. The
sustained release pharmaceutical formulation preferably comprises
gelatin having mesh values between about 8 and 400, but more
preferably between about 18 and 230, and even more preferably
between 35 and 140.
[0093] The reinforcing member may comprise a type of reinforcing
scaffold for dry powdered mixtures or more preferably for powdered
mixtures that have been dispersed into liquid so as to provide
sufficient binding, mechanical support and cohesive integrity
before hydration. When the reinforcing member is a knitted, woven
or non-woven textile, the dry powder mixture or liquid dispersed
mixtures may be dispersed into the interstitial spaces of the
textile. The textile may comprise a fibrous cellulosic material
such as, for example, SafeGauze.RTM. HemoStat.TM. Topical
Hemostatic Dressing commercially available from AMD Medicom, Inc.;
ActCel.TM. Hemostatic Gauze commercially available from Coreva
Health Sciences; SURGICEL.RTM. Original Absorbable Hemostat,
SURGICEL.RTM. FIBRILLAR.TM., SURGICEL.RTM. NU-KNIT.RTM. and
SURGICEL SNoW.TM. commercially available from Ethicon, and others.
The dry powder mixtures can also be reinforced with cellulosic
powders like carboxymethyl cellulose sodium (CMC), SURGICEL.RTM.
Powder Absorbable Hemostat, as well as chopped fibers of CMC or
oxidized regenerated cellulose. The reinforcing members can also be
a made from collagen, alginate, silk, hyaluronic acid, or chitosan,
in the form of a sponge, electrospun felt, porous film or textile.
The dry powder mixtures or liquid dispersed mixtures could be
impregnated into the interstitial spaces of such scaffolds, and the
resulting delivery device could be folded and placed into the tooth
extraction socket, where the delivery device would then be allowed
to hydrate in vivo. However, for reasons pertaining to erosion, the
most desirable approach is to employ a liquid dispersed
mixture.
[0094] When the reinforcing member is in the form of a flexible
textile sheet or scaffold, its geometric shape as well as its
weight percentage in the delivery system can have a significant
effect on the mechanical properties and on the tactile handling
characteristics of the delivery system. Suitable tactile
characteristics have been observed when pharmaceutical formulations
are impregnated into the interstitial spaces of flexible textile
sheets or scaffolds having thicknesses of between 0.01 cm and 0.1
cm, and topical surface areas of between 0.5 cm.sup.2 and 15
cm.sup.2, and more preferably between 1 cm.sup.2 and 9 cm.sup.2,
and even more preferably between 5 cm.sup.2 and 7 cm.sup.2.
Suitable tactile characteristics have also been observed when the
delivery system comprises a cellulose textile as a reinforcing
member at a weight percentage of up to 15% by weight. Moreover,
suitable tactile characteristics have also been observed when the
mass of fiber per topical square centimeter is between 0.005
g/cm.sup.2 to 0.05 g/cm.sup.2, and more preferably between 0.008
g/cm.sup.2 to 0.02 g/cm.sup.2. The mass of fiber per topical square
centimeter is a relative indicator of the bulk density of the
reinforcing member, which can be calculated by dividing the average
weight of the member by its topical surface area. It has also been
found that one or more geometric configurations of the reinforcing
member can be used alone or in combination to form the
formulation-impregnated delivery system. In addition, depending on
the geometric shape of the one or more members, the flexible
textiles can be impregnated and folded in various ways to yield
multilayered impregnated composite structures so that the final
geometric shape of the delivery system is conducive to deployment
by a clinician during end use. In a tooth extraction socket
application, multiple geometric configurations of the delivery
system are suitable so long as the tactile handling characteristics
are acceptable, and as long as the delivery system can be folded,
inserted, and conformed to the shape of a tooth extraction socket,
and provided that the tooth extraction socket is adequately filled
with the delivery system after deployment.
[0095] In order to maximize the amount of anesthetic or analgesic
available for sustained delivery, there exists a need to
simultaneously address the volume-restriction limitations presented
by the size of the wound being treated and ensure that the device
has enough mechanical integrity and cohesive strength to adhere to
the wound mitigate erosion. The reinforcing scaffold for a dry
powdered mixture provides sufficient binding and mechanical support
(i.e., cohesive integrity) before hydration. One could disperse the
dry powdered mixtures of the previous embodiments into the
interstitial spaces of a soft knitted, woven or non-woven textile
such as a fibrous cellulosic material (e.g., SafeGauze,
SURGICEL.RTM. Original, FIBRILLAR, NU-KNIT and SNoW). Conceivably,
dry powder mixtures could be impregnated into the interstitial
spaces of such textiles, and the resulting device could be folded
and placed into the tooth extraction socket, where the device would
then be allowed to hydrate in vivo. However, even with this
approach, the dry powders, although interstitially limited in their
mobility, may still have the propensity to erode and to prematurely
migrate before hydration. Thus, there exists a need to create a
binder system that simultaneously binds the powdered mixtures
together both before hydration, and after hydration, while
simultaneously serving to minimize pre-hydration erosion potential.
Ideally, such a binder system should be capable of being used to
deliver active ingredients for pain management whether it is used
alone, or whether it is used together with a reinforcing member
such as a cellulosic textile. When used with a reinforcing member
like a cellulose textile, the binder system should be compliant
enough to allow for interstitial impregnation, through a process
that minimizes potential damage to the PLGA microspheres (e.g.,
pressing at near ambient temperatures). Once impregnated, the
resulting cellulosic composite should be compliant enough to be
easily folded for placement into an oral tooth extraction socket or
wound, and the tactile feel of the material (i.e., stiffness and
compliance) should be sufficient so as to minimize the potential
for discomfort by the patient.
[0096] It would also be desirable for the non-hydrated binder
system to be optionally useful alone without the use of a
reinforcing textile. In such cases, the binder system could be
allowed to hydrate in vivo, or it could be pre-hydrated and
masticated before insertion into the tooth extraction socket. If
the binder is allowed to hydrate in vivo, it must retain enough
mechanical integrity to resist erosion until it hydrates with
fluids in the tooth extraction socket. On the other hand, the
non-hydrated binder, when impregnated into a reinforcing member
(i.e., a cellulosic textile), would resist erosion to a greater
degree than a non-reinforced binder system, and thus may be a
preferable alternative for in vivo hydration.
[0097] Thus, the sustained release pharmaceutical formulation
comprising a network-forming material optionally impregnates
interstitial spaces of the reinforcing agent, such as a knitted,
woven or non-woven fibrous material, for example, a cellulosic
material like SafeGauze or Surgicel Original. A fibrous textile can
be fit into a tooth extraction socket, wherein the textile is
impregnated with a highly compliant formulation to the degree
permitted by the volume restriction associated with the end use
application. This device takes advantage of the macroscopic free
volume that exists within the interstitial spaces of the textile
and the mechanical reinforcing capability of the textile.
Importantly, mechanical reinforcement enables the use of
mechanically weaker binder formulations that would otherwise be
difficult if not impossible to handle with a pre-hydrated powdered
mixture approach. Highly compliant and mechanically weaker formulas
can equate to the use of lower binder levels and higher microsphere
concentrations to achieve higher bupivacaine dosages. Highly
compliant network forming materials would also be conducive to
simple industrial manufacturing methods for filling the
interstitial spaces of the textile without damaging the PLGA
microspheres, such as continuous pressing under near-ambient
conditions while using the textile as a moving web.
[0098] In addition, if the fibrous material is chosen from a group
of materials with known hemostatic properties, then improved
hemostatic properties can be simultaneously and synergistically
imparted to the delivery device, making it thereby possible for the
delivery device to simultaneously satisfy two additional needs, in
addition to minimizing prep time and to expanding the upper limit
of drug deployment dosages for controlled time-release. First, a
hemostatic fibrous member can impart characteristics that allow the
pharmaceutical formulation to perform the function of a hemostat
during deployment, which can help to facilitate and thereby satisfy
the clinical need for clot-formation and protective scab formation.
Secondly, the fibrous reinforcement can continue to facilitate the
formation of a mechanically stable, compliant, persistent, and
erosion-resistant scaffold-like composite that resists dislodging
during use by simultaneously interacting with cavity fluids such as
saliva and blood and with formula ingredients as they inter-diffuse
and mix together under static conditions over time. This function
would help protect the resulting scab from dislodging and would
thereby help to prevent the painful occurrence of dry socket, a
very important clinical need.
[0099] Thus, the use of hemostatic fibrous reinforcement material
in the delivery device simultaneously provides many desirable
features. The fibrous reinforcement facilitates initial composite
reinforcement of the pharmaceutical formulation during
manufacturing, during storage, and during initial deployment. The
fibrous reinforcement allows for the optional use of lower
network-forming material levels in the formula thereby expanding
the upper limit for dispersed active ingredient and drug dosage,
and for the optional use of lower levels of higher molecular weight
network-forming materials in the binder phase of the formulation
thereby providing reduced viscosity for ease of manufacturing and
for higher initial compliance for handling efficacy. The fibrous
reinforcement provides the advantage of hemostatic properties and
simultaneous composite reinforcement during initial deployment into
the socket and, if the fiber reinforcing member is properly chosen,
the fiber reinforcing member can also continue to reinforce the
composite during extended periods under static conditions after
deployment, thereby facilitating in vivo composite formation with
fluids in the socket while minimizing the propensity for erosion.
This can facilitate formation of an in vivo composite that not only
protects the forming scab from premature dislodging, but provides a
vessel for the formulation to persist and to continue to perform
its drug delivery function over prolonged periods without being
prematurely ejected or eroded from the tooth extraction socket.
[0100] It can be appreciated by those skilled in the art of
composite materials that the physical properties and
handling-related characteristics of composites like those described
herein can be influenced by many fiber-related factors including,
for example, the density of individual fibers and fiber bundles;
the density of knitted, woven or non-woven textiles comprising
fibers and fiber bundles; the bulk density of the fibrous members
whether they are knitted, woven or non-woven; the geometric length
of fibers and fiber bundles; the total surface area per unit weight
of the fibrous members; the surface wetting characteristics of the
fibrous members towards both hydrophobic and hydrophilic materials;
the volume and weight ratios of the fibrous members to the formula
members; and among other factors, the rate of dissolution of the
fibers in vivo, as influenced by their solubility, their degree of
oxidation, their molecular weight, and their surface wetting
characteristics. Each of these fiber-related factors, either alone
or in any combination, can have a profound impact on the composite
device's manufacturability, on its mechanical properties during
initial deployment, and on its dynamically evolving properties as
the device experiences static diffusion and intermixing with tooth
extraction socket fluids during the entire timeframe associated
with the near static condition of the in vivo environment,
particularly during the entire end use period associated with the
wound healing cycle and with the drug delivery.
[0101] As one aspect of this invention, it can be appreciated that
the choice of the fibrous member for the composite device is an
important one, and that the material can be tuned to the
application by controlling the degree of oxidation which affects
solubility, the molecular weight of the cellulose, the fiber
surface area per unit volume, the fiber bundle density, the bulk
knit density, and the like. Aside from these tunable factors, it is
also possible to use a mixture of fibrous member types. For
example, the fibrous composite could be comprised of both a
relatively fast-dissolving type of fiber member (e.g., SafeGauze),
and a relatively slow-dissolving member (e.g., Surgicel Original).
Use of multiple fiber types can impart combinations of desirable
characteristics, including faster initial wetting and better
initial adhesion during deployment from the more soluble fiber
member, and longer-term composite integrity during the in vivo use
period associated with dynamic changes in properties owing to
inter-diffusion of tooth extraction socket fluids with the device
from the less soluble fiber member.
[0102] In one aspect, pH modulators may be used as a component to
adjust the pH of the formulation. Bases or buffering additives,
such as di-sodium citrate, and acidic additives, such as ascorbic
acid or citric acid, can be provided at selected levels. Initial
gelation rates and viscosities of gelatin can be modulated by
protonation, for example, with citric acid. Protein-moiety
protonation induces faster gelation and higher relative
viscosities. Thus, slower or faster gelation rates can be achieved
by modulating pH as desired. As such, gelatins can be used for
formulating PLGA microsphere-containing formulations with
mechanical and gelation characteristics that vary depending upon
gelatin-type and pH.
[0103] It is important to recognize that the efficacy of the device
will be impacted by the diffusion rate of active ingredients, such
as bupivacaine. This diffusion rate will not only be affected by
microencapsulation of bupivacaine with PLGA, it will also be
affected by the water-solubility of bupivacaine, which is affected
by the equilibrium concentration of bupivacaine's protonated
acidic-form in competition with its non-protonated free-base form.
In the presence of Bronsted acids (e.g., protons from citric acid,
protons from protonated amine moieties from the gelatin protein,
etc.), the free-base form of bupivacaine and drugs with similar
chemical structures will protonate to some degree, and the more
water-soluble protonated form will exist in equilibrium with the
less water-soluble free-base form. To this end, there can be an
advantage associated with using pH-modulators like citric acid to
assist in controlling the effective solubility of bupivacaine.
[0104] The relative concentrations of protonated and non-protonated
bupivacaine structures will be affected by all competitive
acid-base reactions, including those involving protein amine
moieties. For example, given that different proteins will exhibit
differing degrees of acid neutralization capacity, and given that
the relative viscosities can increase in the presence of a proton
source (this is demonstrated in Example 1), it follows that free
base drug diffusion rates will differ in the presence of different
protein-types (i.e., for reasons pertaining to drug solubility, and
for reasons pertaining to diffusion rates being attenuated by
increased viscosity). PLGA hydrolysis rates will also be affected
by pH, and by competitive equilibrium reactions with other Bronsted
bases (e.g., di-sodium citrate, protein amines, and the free-base
form of bupivacaine).
[0105] Thus, if one were to add an acid such as citric acid to a
pharmaceutical formulation with the intent of skewing the
bupivacaine acid-base equilibrium towards the more water-soluble
protonated form, the relative equilibrium concentration of the more
soluble protonated form would vary depending on the composition of
the chemical environment. For example, a chemical environment
comprised of different types of Bronsted bases (e.g., protein
amines from various types of gelatins), and different types of
Bronsted acids (e.g., citric acid, ascorbic acid, sebacic acid,
etc.) would lead to different equilibrium concentrations of the
more water-soluble, protonated form of bupivacaine. Accordingly,
apparent drug activity and release rates would be affected for this
reason. Similarly, if the hydrochloride salt of bupivacaine (i.e.,
the more water-soluble protonated form) were added to a formulation
with these types of gelatins under pH neutral conditions (i.e.,
with no additional acid or base), the protonated bupivacaine would
enter into equilibrium with competitive Bronsted bases from the
protein gelatins whereby the acid neutralization capacity of the
protein would govern the ultimate equilibrium concentration of the
more water-soluble protonated form of the drug.
[0106] Importantly, the pH of the chemical environment will also
have an impact on the rheological characteristics of the
formulation. This in turn will not only have an impact on the
diffusion rate of active molecules like bupivacaine, but it will
also have an impact on the compliance characteristics of the
formulation, which in turn will affect its formability within a
fixed volume cavity, such as a tooth extraction socket.
[0107] The effects of pH on properties of the network-forming
material can be appreciated by one of ordinary skill in the art.
The ratio of citric acid or other alternative acids to protein can
be used to achieve a gel network with suitable mechanical
integrity. Higher citric acid to gelatin ratios would lead to even
faster gelation rates as would lower levels of water.
[0108] For rheological purposes, a pharmaceutical formulation
comprising powdered mixtures may be developed with commercial
gelatins, whereby a formulation will optionally incorporate a level
of citric acid or other types of acids, which when mixed together
with an appropriate ratio of water to solids will allow for a rate
of gelation that is desirable for clinical use. Further, the
formula composition can be controlled so as to exhibit appropriate
elastic modulus and compliance characteristics by modulating
factors such as the water level, the molecular weight distribution
of the gelatin (Mw/Mn), the concentration of acid, and the type of
acid additive, etc. Simultaneously, it is understood that a balance
would be achieved with other factors that impact the efficacy of
the pharmaceutical formulation, including the aforementioned
chemical environment factors that affect the solubility and
diffusion of bupivacaine, including the initial concentration of
bupivacaine in its free-base and in its acidic form and the
ultimate equilibrium concentration of both species within the end
use chemical environment.
[0109] In one aspect, the pharmaceutical formulation delivers a
maximum level dosage of bupivacaine (BUP) into a fixed volume
cavity, assuming an occupied formula volume of 1 cc for the oral
post tooth-extraction cavity. The target dosage range for
bupivacaine is between a level approaching possible toxicity on the
high-delivery side and a level representing clinical usefulness on
the low-delivery side. The formulation composition is dependent on
the bupivacaine target dosage level due to the unique occupied
volume limitation for this type of end use application, with a
maximum bupivacaine dose estimate targeted to be up to 360 mg over
a 4-day period (90 mg/day.times.4). Three different pathways were
identified to address the problem of maximizing dosage: (1)
increasing bupivacaine-loading to its maximum theoretical level of
about 50% w/w within the PLGA microspheres; (2) minimizing the
network-forming material levels to the extent permitted without
simultaneously deteriorating mechanical properties; and (3)
minimizing the level of water required for hydration/mastication to
the extent tolerable without experiencing unmanageable decreases in
compliance.
[0110] In an embodiment, dry powders of bupivacaine-loaded PLGA
microspheres and gelatin would be masticated with water to be
delivered as a compliant dough-like material in end use for a
desirable bupivacaine release profile. Volume restriction for the
application estimated to be ca. 0.55 cc causes the dosage of
bupivacaine to be severely limited by the occupied volume fraction
of binder and water. Higher levels of bupivacaine loading in the
PLGA microspheres are desirable for achieving useful bupivacaine
delivery dosages, higher than the 20% w/w level that was used in
the prior art since this level would only lead to maximum dose
deliveries of less than 60 mg. Lower binder levels are required to
maximize the microsphere content and hence the bupivacaine delivery
dosage, which is a constraint that weakens the composite and
necessitates not only better network-forming binders, but higher
levels of volume-occupying water for plasticization. Lower binder
levels necessitate higher molecular weight network-forming gels
that are susceptible to time-dependent reductions in compliance
owing to diffusion-rate limitations which impact the time required
for the network to reach its equilibrium state. Diffusion rates and
time-dependent compliance behavior are further confounded by both
the particle size distribution of the microspheres, which affects
the bupivacaine time-release profile, and by the size of gelatin
particulates. From a mechanical property perspective, it is
desirable to maximize the smaller particle size fraction while
simultaneously balancing the overall distribution to achieve the
desired bupivacaine release profile since smaller particles will
release faster than larger ones.
[0111] One embodiment of the sustained release pharmaceutical
formulation using powder mixtures appears to deliver a dosage of
about 140 mg bupivacaine to a 0.55 cc cavity, and only then by
assuming that the % bupivacaine loading in the microspheres is
increased from 20% to 50% by weight. Low gelatin binder levels are
also required to maximize the volume fraction of microspheres and
bupivacaine. It appears that the lower limit threshold for the
network-forming material is approximately 18% of the dry weight. At
these levels, a network-forming gel like bovine gelatin is
preferred as having sufficient strength to bind the spheres
together. If the product is intended to be masticated with water,
and if higher bupivacaine dosages are desired, then the occupied
volume of water must also be accounted for, and the water-level
should be minimized since it will effectively dilute the
microsphere concentration and will further reduce the maximum
bupivacaine delivery dosage. For reasons pertaining to mechanical
properties, it is also preferable to skew the PLGA particle size
distribution towards small particles to the degree that this can be
tolerated depending on bupivacaine release profile targets. Larger
microspheres made via an emulsion process provide qualitatively
lower formula viscosities than their spinning-disc/spray-dried
counterparts. In essence, this equates to a higher PLGA loading
potential during mixing, which is also directionally preferred for
achieving higher bupivacaine dosages; but only to the degree that
adequate compliance and cohesive strength can be maintained. The
D50=42.1 micron emulsion particles were also observed to mix more
uniformly with faster wetting than their similarly-sized
spinning-disc spray-dried counterparts, the D50=42.7 micron placebo
PLGA microspheres. Again, smaller particles, D50=3.4 microns as
made by spinning-disc methods, by spray-drying, or by
emulsion-solvent extraction processes, are desirable for reasons
pertaining to mechanical properties, but only to the degree that
their higher surface-to-volume ratios and release characteristics
can be conducive to achieving specific time-dependent bupivacaine
release profile targets.
[0112] Although release profile targets will be end use specific,
it should be understood that there will be several adjustable
factors besides PLGA surface-to-volume ratios that can also
conceivably be used to modulate and control the time-release
profiles of bupivacaine and the like. For example, citric acid (a
Bronsted acid) or di-sodium citrate (a Bronsted base) was observed
to be viable with no obvious deleterious effects on rheology or
properties of the delivery system. Citric acid was observed to
enhance binder system network formation of the network-forming
material. From this perspective, these types of compounds can serve
dual functions: not only can they be used to modulate the physical
properties of the binder system, their activity can also be
exploited for the dual purpose of modulating the solubility of the
bupivacaine free base. For example, a Bronsted acid will enhance
the solubility of bupivacaine free base as it is released from a
PLGA particle, thereby enhancing its bioavailability. Conversely, a
Bronsted base would skew the acid-base equilibrium towards more
bupivacaine free base, thereby reducing its bioavailability.
Further, these types of compounds can be employed directly as
powdered ingredients, which would make them immediately available
upon device hydration. In addition, these types of compounds can be
optionally and separately microencapsulated, which would attenuate
their availability for acid-base interactions with bupivacaine in
its acidic form or its free-base form. By balancing these types of
formulation levers, it can be appreciated that one could achieve
targeted bupivacaine release profiles while simultaneously
employing higher fractions of high surface-to-volume particles if
so desired. For example, with the combined use of these levers, one
could potentially use a higher fraction of 3.4 um particles than
would otherwise be viable. Again, this direction might be desirable
for reasons pertaining to achieving improved mechanical properties,
which in turn could be leveraged to achieve lower net binder
levels, including the network-forming material, and higher net PLGA
levels and accompanying higher net bupivacaine dosages if so
desired.
[0113] As described above, hydrophobic components in the
pharmaceutical formulation and related delivery devices can be
desirable from the standpoint that they can be formulated to yield
dough-like materials with compliance characteristics that are
conducive to end use deployment, without having to rely upon
pre-deployment swelling and gelation of the gelatin particulates.
Thus, the pharmaceutical formulation comprising hydrophobic
components and related delivery devices are ones whereby the
gelatin particulates remain intact during manufacture and storage,
and do not yield macroscopic chain-entangled gelled networks until
they become exposed to a tooth extraction socket and its fluids
after deployment unless the option of pre-deployment hydration is
exercised.
[0114] It is important to note that each of the embodiments of the
present formulation will eventually become hydrated with fluids
from the tooth extraction socket after deployment. This is
predominantly due to the presence of the hygroscopic,
water-absorbing network-forming polymer, like gelatin, or to the
presence of other water-absorbing materials such as cellulose
fibers. However, in order to render the devices as compliant and
conformable prior to their deployment, it is desirable that they be
properly formulated in advance of deployment, so that the clinician
does not have to spend time meticulously measuring and premixing
materials before they can be used. In other words, it is desirable
to have a device that is already a compliant solid without having
to be premixed with fluids like saline solutions or water.
[0115] In an embodiment of the present formulation comprising at
least one hydrophilic component, water may be used as a plasticizer
to hydrate and to masticate blends of the powder ingredients to
yield a compliant dough-like mixture (e.g., water+bovine gelatin
with PLGA-encapsulated BUP as described in Example 12). In these
cases, water is the primary liquid ingredient in the pharmaceutical
formulation, and the mechanical integrity of the formulation is
achieved by virtue of gelation and network formation prior to the
deployment. The compliance and conformability of this formulation
is controlled by the ratio of water to gelatin (w/w) with
consideration also given to the total % solids in the plasticized
mixture. Importantly, water is used as a liquid plasticizer for the
gelatin polymer in this embodiment. A plasticizer is generally a
liquid (sometimes a solid) that when blended with a polymer,
increases the fraction of free volume, which in turn lowers the
polymer glass transition temperature, and consequently lowers the
elastic modulus, and increases the compliance. Plasticizers are
known to be at least partially miscible with the polymers that they
plasticize.
[0116] In an embodiment of the present formulation comprising at
least one hydrophobic component, oils with optional waxes are used
as liquid carriers to suspend hygroscopic, water-absorbing
network-forming materials, such as gelatin powders together with
other dispersed ingredients, including PLGA-encapsulated BUP, free
(non-encapsulated) BUP, and citric acid, to name a few. This
embodiment achieves pre-deployment conformability and compliance
characteristics not by plasticization of a polymeric continuous
phase, but instead by virtue of other interactive factors that
impact the rheological properties of suspensions, including the
ratio of hydrophobic liquid to wax, which controls the viscosity of
the liquid carrier and affects the viscosity of the resulting
vehicle, the particle size distributions of dispersed ingredients,
and the total percentage of dispersed solids, to name a few. In
these cases, the mechanical integrity of the pre-deployed
formulation is not achieved by virtue of gelling a polymer with a
plasticizer to yield a reinforcing polymer network, but instead by
virtue of fiber reinforcement by impregnating knitted or woven
cellulose textiles or non-woven fibers with non-gelled suspensions
to yield fiber-reinforced composite-like structures.
[0117] Thus, one of the primary distinctions between the
hydrophilic and hydrophobic formulations and delivery systems
relates to pre-deployment morphology. By design, a hydrophilic
formulation or delivery system is comprised of a water-miscible
hygroscopic polymer network that is homogenously gelled and
pre-plasticized with a liquid such as water, glycerin, honey,
polyethylene glycols, polypropylene glycols, etc. By contrast, the
hydrophobic formulation or delivery system contains inter-dispersed
suspended particulates of water-miscible and hygroscopic
network-forming polymers like gelatin that have the latent
potential to form gelled networks once exposed to water after
deployment, but in their pre-deployment state, they are made to
persist as morphologically discrete entities suspended within and
wetted by a hydrophobic liquid. These hydrophobic formulations and
delivery systems (sometimes interchangeably referred to as devices
herein) do not rely on gelatin plasticization and network formation
to achieve their pre-deployment properties. However, after
deployment, they are morphologically designed to accept water
through diffusion of oral fluids like saliva and blood, which
allows for post-deployment polymer network formation, analogous to
what occurs in the pre-deployment stage with a hydrophilic
formulation or device. At that point after the deployment, the
development of a gelled polymer network from water-ingress can have
the added benefit of providing an additional mechanism of
mechanical reinforcement, augmenting that which may already be
provided by inter-dispersed cellulose fibers.
[0118] With these morphological considerations in mind, the
differences between a hydrophilic and hydrophobic embodiment of the
pharmaceutical formulation can be further reduced to another
important design-controlling distinction, namely, the nature of the
liquid component that is used in formulation. Generally, a liquid
that leads to pre-deployment gelation of the network forming
component is best suited and preferred for use in preparing the
hydrophilic embodiment of the formulation. A liquid that does not
lead to pre-deployment gelation of the network forming component,
or at least little to no gelation for a period of time after
manufacture that coincides with the desired shelf-life prior to
deployment, is best suited and preferred for use in preparing the
hydrophobic embodiment of the formulation. The delineation between
a liquid that leads to gelation of the network forming component
and one that does not lead to gelation can be defined by a
suspension test as demonstrated Example 14. In general, if there
are no signs of gelation within a predetermined monitoring time
window, then the liquids are considered to be candidates for use as
a "hydrophobic" component of the formulation. Mineral oil, caprylic
triglyceride, isopropyl palmitate, and coconut oil are such
liquids. Liquids that are observed to lead to gelation of gelatin
within the time monitoring window are considered to be good
candidates for use as a "hydrophilic" component of the formulation.
Glycerin and water are such liquids. Note that similar tests can be
employed to test the miscibility of carrier liquids with other
dispersed ingredients.
[0119] In some circumstances, the degree of hydrophilicity and
hydrophobicity of a liquid can also be gauged by parameters that
pertain to molecular-level properties such as polarity (e.g.,
dipole moment forces from permanent dipoles), dispersion forces
(e.g., non-permanent dipoles or van der Walls forces), and hydrogen
bonding forces. Indices such as the Hildebrand Solubility Parameter
(HSP) or Hansen Solubility Parameter (HAN) of liquids and polymers
(J. Brandrup and E. H. Immergut, Polymer Handbook, Third Edition,
John Wiley & Sons, New York, 1989, pp. 519-559), as well as Hoy
solubility parameters (HOY), have been developed in attempts to
better quantify what is meant by "hydrophilicity" and
"hydrophobicity." Hoy solubility parameters (HOY), like Hansen
Solubility parameters (HAN) are based on chemical group methods of
calculating energetic contributions from dispersion forces, polar
forces, and hydrogen bonding forces. These contributions are summed
to yield the total solubility parameter by taking the square root
of the sum of the squares. Generally, although the estimation
methods differ for the HAN and HOY terms, the sums of the
contributions from HAN and HOY parameters produce similar total
solubility parameter estimates, which are also considered to be
equivalent to HSP values (i.e.,
HSP.about.HAN.sub.total.about.HOY.sub.total).
[0120] It is generally understood by those skilled in the art that
polymers and liquids tend to be more miscible when their solubility
parameters are similar in magnitude to one another, as the
differences between them approach zero. Conversely, polymer/solvent
pairs become less miscible as their solubility parameters diverge
from one another, as the differences between them become
greater.
[0121] For the purposes of the present invention, the most
hydrophobic liquids can be defined as those with either a small or
no permanent dipole moment, and with a low capacity to participate
in hydrogen bonding. These types of liquids have been observed to
be the least compatible with highly polar and water-soluble
protein-based polymers like gelatin, which explains why the gelatin
particulates remain dispersed and stable over time when suspended
(i.e., not gelled) in formulations comprised of such liquid
carriers. These types of liquids would also be expected to have
limited compatibility with other polar molecules, such as water and
BUP-HCl, thus rendering them as relative deterrents against both
molecular-level and macro-level diffusion during the end use
application. This behavior renders such liquids as useful levers in
achieving specific control over time-release profiles. In the
present formulation, an example of this type of liquid is
represented by a paraffinic hydrocarbon like mineral oil.
[0122] On the other side of the spectrum, liquids with permanent
dipoles and with higher capacities for hydrogen bonding can be
classified as being less hydrophobic and more hydrophilic. In the
present pharmaceutical formulation, this type of liquid is
represented by water in one extreme (HSP=approximately 48
MPa.sup.1/2). These types of liquids are highly compatible with
hygroscopic polymers like gelatin, which explains why the dispersed
gelatin particulates do not persist in formulations containing
water, but instead become swollen through diffusion and
plasticization, leading to the coalescence of the particulates
through polymer chain entanglement, and leading ultimately to
gelation and to solid network formation prior to deployment of the
pharmaceutical formulation.
[0123] Note that for the case of a pharmaceutical formulation that
is prepared with hydrophobic components like oils or waxes, the
more hygroscopic components, like gelatin particles and cellulose
fibers, remain discrete and intact prior to hydration, either as
dispersed, non-gelled particulates, or as intermeshed fibrous
entities. In these cases, the oils and waxes that constitute a
continuous phase of the pharmaceutical formulation serve to
facilitate the dispersion of other ingredients like gelatin, PLGA
microparticles, BUP, and citric acid. Note that optional
surfactants can also be added to assist in stabilizing such
dispersions.
[0124] In its pre-deployment morphological state, the mechanical
integrity of the pharmaceutical formulation comprising a
hydrophobic component may be derived from its reinforcement with
cellulose fibers. Importantly, the morphology of the hydrophobic
formulation has been designed to adsorb polar liquids like water.
Thus, when a polar liquid (e.g., water, glycerin, polyethylene
glycol, mixtures thereof, or fluids from the tooth extraction
socket, etc.) is intermixed with a formulation having the
hydrophobic component, the morphology of the formulation
accommodates the adsorption of the polar liquid without producing
the side-effect of macroscopic phase separation of other
components. This behavior is consistent with a morphological change
that occurs when polar liquids are mixed with the formulation,
whereby the more hygroscopic components like gelatin or cellulose
begin to absorb the polar liquid, becoming plasticized, and then
begin to coalesce into a gelled network matrix such that the new
continuous phase contains the gelled network matrix, including
polar liquid+gelatin+cellulose, inter-dispersed together with the
hydrophobic components, the oils and waxes that previously
constituted the continuous phase prior to hydration. At this stage,
other dispersed ingredients like PLGA, BUP, BUP-HCl, citric acid,
etc., that were previously dispersed in the oil-based continuous
phase, either remain dispersed within the oil-phase components that
themselves become inter-dispersed within the gelled matrix, or they
become directly dissolved in the water that diffuses into the
newly-formed continuous phase of the gelled matrix. Importantly,
the plasticization, the chain-entanglement, the ensuing gelation,
and the ultimate network formation that accompanies this adsorption
process are desirable attributes for the hydrophobic pharmaceutical
formulation. Most importantly, and by design, this morphological
change is made to occur in vivo and does not have to occur during
the pre-deployment stage or during the storage period.
[0125] The latent capacity for a pharmaceutical formulation
comprising a hydrophobic component to adsorb a polar H-bonding
liquid like water is not only a desirable and surprising attribute
that arises from the synergistic interactions among the component
ingredients of the formulation, it is a measurable attribute that
can be used to specify a distinguishing characteristic of a
hydrophobic pharmaceutical formulation. Namely, a hydrophobic
pharmaceutical formulation, as used herein, is one that after being
mixed via physical mastication with water at a minimum ratio of
water to device=0.2/1 w/w, or more preferably 0.33/1 w/w, or even
more preferably 0.44/1 w/w or higher, does not exhibit macroscopic
phase separation under static conditions for a period of at least 1
hour, and preferably for 2 or more hours, and more preferably for
24 hours. The formulation further retains the added water for said
period of time under static conditions without exhibiting visual
indications of macro phase separation of water or other
components.
[0126] As stated previously, if the end-product objective is to
minimize active-ingredient dilution in the pharmaceutical
formulation while simultaneously achieving mechanical compliance
characteristics that are desirable for deployment, then gelation of
gelatin or other macromolecular hygroscopic network-forming
components would be most desirable if it were made to occur after
deployment and not before. Thus, the pharmaceutical formulation
comprising a hydrophobic liquid like mineral oil or others as shown
in Table 14-3 represents a preferred approach towards achieving
this objective.
[0127] On the other hand, when compared to hydrophobic liquids like
mineral oil, hydrophilic liquids like water and glycerin are more
compatible and more miscible with polar molecules like BUP-HCl, a
fact which is consistent with the observation of faster active
ingredient diffusion rates exhibited by the pharmaceutical
formulation pre-plasticized with water as opposed to those prepared
with mineral oil as the liquid vehicle carrier. Hence, if the
end-product objective is to maximize the release rates of
water-soluble active-ingredients while simultaneously achieving
mechanical compliance characteristics that are desirable for
deployment, then gelation of gelatin or other hygroscopic
network-forming components with hydrophilic liquids like water and
glycerin could be a desirable approach before deployment. Thus, the
pharmaceutical formulation comprising a hydrophilic component
represents a method of approach towards achieving this objective,
but only if the resulting dilution of active ingredients can be
tolerated in the end use application.
[0128] Again, in the absence of gelation, the hydrophobic
pharmaceutical formulation achieve their initial mechanical
cohesive integrity through a mechanism that is independent of
gelled network formation. Specifically, if the formulation is to
have the compliance characteristics of a cream, it can then be used
to disperse active ingredients, and it can then be impregnated into
a fibrous textile which serves as a reinforcing scaffold forming a
delivery device before its deployment. The reinforced device is
therefore made to have cohesive integrity and compliance which
renders it as sufficiently acceptable for use by the clinician
during its deployment. It is only later, after deployment, that the
dispersed gelatin particulates and wetted and impregnated cellulose
fibers begin to swell with liquids from the tooth extraction
socket, leading to their chain entanglement, and ultimately to
their network formation and to an accompanying change in
morphology. The gelled network then becomes a type of reinforcing
scaffold for the device in vivo, serving to enhance the cohesive
strength of the pharmaceutical formulation, which enhances its
mechanical integrity after deployment, and not before.
[0129] Other liquids besides mineral oil, such as caprylic
triglyceride and isopropyl palmitate, are more polar than mineral
oil, and they have at least some capacity for hydrogen bonding.
However, their polarity and H-bonding characteristics are
insufficient to cause gelation of the gelatin particulates that are
suspended within them. Thus, although these types of liquids have
permanent dipoles and therefore have some capacity for hydrogen
bonding, they are poor plasticizers for gelatin. For the purposes
of description, a pharmaceutical formulation or delivery device
comprised of such liquids may be referred to as "hydrophobic". The
"hydrophobic" formulation containing them have a distinguishing
attribute in common: the liquid carriers that serve to suspend and
bind the ingredients do not promote the gelation of the gelatin
particulates, and they are either immiscible with gelatin or have
limited miscibility under ambient conditions. Consequently,
macromolecular chain entanglement and gelation do not occur when
the particulates are suspended in such liquids.
[0130] However, liquids that are deemed as being suitable for use
as a hydrophobic ingredient can also perform other functions when
formulated into the pharmaceutical formulation. For example, the
HAN of isopropyl palmitate is reported as 15.3 MPa.sup.1/2.
Although these types of liquids are recognized as being more polar
than mineral oil, for the purposes of the formulation, they are
still classified as being relatively hydrophobic in that they do
not diffuse and swell gelatin particulates in the way that water
does. Instead, the gelatin protein particulates persist in such
formulations until they are subjected to hydration during end use.
Nevertheless, the permanent dipole moments of these liquids would
be anticipated to render them as more amenable to facilitating
molecular-scale diffusion of small polar molecules than would
mineral oil. Thus, liquids of these types can be useful to modulate
diffusion rates of active ingredients, thereby providing an
additional lever to achieve intermediate controlled-release time
profiles. In addition, hydrophobic liquids with higher polarity
than mineral oil can also serve the secondary purpose of lowering
the Tg of PLGA via plasticization. This would result in a faster
rate of diffusion of encapsulated ingredients because a lower Tg
will equate to a higher fraction of free volume, which in turn
would translate to lower potential energy barriers for diffusion of
small molecules across the PLGA polymer gradient from within the
PLGA particle and into the matrix.
[0131] There are occasions when the use of a pharmaceutical
formulation comprising a hydrophilic component gelled with a
hydrophilic liquid before deployment would be desirable for end
use. For example, a hydrophilic formulation that is mixed with
water can be useful in achieving relatively fast time-release
profiles of water-soluble ingredients. This embodiment of the
pharmaceutical formulation is first premixed and pre-plasticized
with water, glycerin, polyethylene glycols, other polyhydric
alcohols, or mixtures thereof. This embodiment is analogous to the
hydrophobic embodiment, but they are made with a polar H-bonding
liquid as the primary liquid ingredient instead of oils and waxes,
and they are designed to gel prior to deployment instead of
afterwards. Thus, as long as they are shelf-stable against
microbial growth, these types of pre-gelled formulations can be
used as control-release delivery devices on their own--without
fiber reinforcement. However, they can also be optionally
reinforced with a fibrous cellulose hemostat to form a composite
structure.
[0132] As noted by Jaymin C. Shah and Manoj Maniar in Journal of
Controlled Release, 23 (1993) 261-270, controlled release of active
ingredients like BUP from polymeric matrices can occur via
diffusion, dissolution or erosion of the polymer. The authors note
that erosion or diffusion processes are generally assumed to
control the rate of drug release. Hence, if the drug and its
conjugate salt have low water solubility, then it is anticipated
that the dissolution rate of the drug could have a significant
effect on the release-kinetics of the drug.
[0133] It should also be realized that diffusion and erosion are
interactive processes, and that diffusion involves not just the
egress of active ingredients from the formulation, but ingress of
water and fluids from the chemical environment where the
formulation is deployed. As fluids diffuse into the formulation via
both macro and molecular-level pathways, the matrix polymer can
become more susceptible to erosion, either through dissolution of
volume elements from the exposed surfaces, from the macro
separation of particulates near the surfaces, or through a
combination of the two.
[0134] As noted earlier, one advantage of using fibrous
reinforcement for a delivery device is that it can improve the
cohesive integrity of the pharmaceutical formulation, and thereby
render it to be more erosion resistant. When a formulation erodes
during end use, internal cohesive failures of the matrix can cause
particulates of the device to become macroscopically separated from
the original structure. During end use, fluids can permeate into
the matrix phase through a combination of macroscopic and
microscopic diffusion mechanisms. Macroscopic diffusion can occur
through permeable boundaries that are present from defects like
void elements arising from entrapped air between partially bonded
matrix polymer particulates (e.g., gelatin particulates), or from
matrix polymer that is partially delaminated from the surfaces of
weakly bonded elements or components that are dispersed within the
matrix.
[0135] If the pharmaceutical formulation comprises a polymer that
is hygroscopic, molecular level diffusion of hydrous liquids can
occur along every frontal boundary that becomes available to the
fluid. When the fluid macroscopically diffuses into the matrix
along a frontal boundary, it also can begin to permeate into the
matrix polymer through a process of molecular-level diffusion. As a
volume element of a matrix polymer begins to expand from the
ingress of lower molecular weight fluids, it can become plasticized
by the fluid, leading to an increase in the fraction of free volume
within the matrix polymer phase, and to a subsequent further
increase in the rate of molecular level diffusion, both into and
out of the matrix polymer network.
[0136] An increase in free volume at the molecular level also leads
to a number of additional physical changes in the matrix polymer
phase, including a decrease in the glass transition temperature, an
accompanying decrease in modulus, a decrease in ultimate stress to
failure (i.e., lower strength), and to an accompanying acceleration
in the rate of molecular level diffusion of molecules both into and
out of the matrix polymer phase. The macro volume expansion of the
liquid-occupied volume element, which is the polymer volume element
that has become diffusion-permeated and plasticized by fluids,
leads to the development of localized stresses that tend to
accumulate at weak boundaries, which are at frontal boundaries that
separate swollen volume elements from other volume elements that
have not yet been permeated and are not yet swollen. Defects sites
near these boundary regions become particularly susceptible to
localized stress-induced tensile and shear types of failures. The
ensuing number of internal cohesive failure events can begin to
increase and even to accelerate from excessive strains at weak
junctures, for example at cell walls of macroscopic voids, at the
interfaces of weakly bonded particulates, etc. The cycle continues
as more macroscopic pathways develop for the macroscopic ingress of
even more fluids, leading to a further increase in the number of
pathways for molecular level diffusion, which then leads to an
increase in the number of swollen volume elements, which then leads
to the further development of more localized stresses. Hence, the
cascade continues, culminating in an acceleration in the rate of
occurrence of ultimate failure events.
[0137] The interconnected processes of erosion and diffusion can
also affect the efficacy of the pharmaceutical formulation.
Clearly, as erosion occurs, the total amount of surface area
simultaneously increases. This will affect one of the primary
functions of the formulation- to achieve and maintain a specific
time-controlled release profile of one or more active ingredients
during end use. An increase in the total surface area from erosion
leads to an acceleration of molecular-scale diffusion of active
ingredients across the growing number of concentration gradients
that are provided by the growing number of interfacial boundaries.
This process will not only impact the molecular level diffusion
rates through the matrix polymer, it can impact the molecular level
diffusion rates through other types of secondary diffusion barriers
that have been purposely put into place, such as the diffusion
barrier created by an encapsulating PLGA polymer which serves to
impede the molecular-level diffusion rate of its encapsulated
active ingredients like BUP or BUP-HCl.
[0138] Any process that leads to an increase in free volume of a
polymer will subsequently lead to an increase in the number of
molecular pathways that are available for molecular level
diffusion. Importantly, diffusion of small molecules will occur
across passive boundaries where a concentration gradient is in
existence (i.e., Fickian diffusion). Aside from relative polarity
considerations, the rate of diffusion depends on the fraction of
free volume within the materials on both sides of the frontal
boundary, as well as the relative concentration of the diffusing
species on both sides. Thus, as fluids begin to have access to the
surfaces of PLGA particles within the formulation, they can
permeate the surfaces of the particles and thereby increase free
volume, and then increase the rate of diffusion of small molecules
that are encapsulated and contained within them. To add even more
complexity to this scenario, if the fluid contains water, PLGA can
hydrolyze. The hydrolysis process leads to a decrease in molecular
weight, to the production of more chain ends, and thus to a further
increase in free volume which further enhances the rate of
diffusion. A gelatin matrix polymer with polypeptide sequences will
also be susceptible to the same type of hydrolysis-initiated
acceleration of free volume. Thus, each molecular level diffusion
barrier that is purposely set in place to control the release of
drugs and the like can become altered and affected by a cascade of
macroscopic and molecular-level events. These events will
collectively affect the global time release profile of the
formulation. Of course, when harnessed for the purpose of achieving
specific control-release profiles over sustained periods of time,
these mechanisms can be useful. On the other hand, if these
processes occur too quickly, it may become difficult if not
impossible to achieve longer-term sustained release.
[0139] Importantly, composite structures can be used to reduce the
rate of occurrence of internal cohesive failure events of the types
described above. In a composite-like structure, the matrix can be
reinforced with fibers or with particulates, which serve as
scaffolds that can help to hold a mechanically weaker matrix phase
in place by reducing the probability of crack growth and
propagation along any one single boundary via distributing stresses
from swelling over larger volume elements and hence over multiple
boundaries within the structure, thereby reducing the magnitudes of
localized stresses and strains, and hence reducing the number and
frequency of catastrophic failure events. Lower levels of localized
stresses will translate to lower localized strains, which in turn,
depending on the geometric structure of the defect site, can lead
to sustained mechanical and cohesive integrity of the delivery
device over longer periods of time.
[0140] The pharmaceutical formulation compromising a hydrophobic
component lends itself well to the creation of fiber-reinforced
composites, primarily because by design, the formulation that is
used to impregnate the fibers are not pre-gelled into macro
polymeric networks. Instead, the formulation is, with their
hydrophobic liquid carriers, remaining compliant and moldable for
long periods of time. The gelatin particulates suspended therein do
not begin to gel and swell until they are exposed to fluids within
the tooth extraction socket. Even then, the rate of water ingress
is diminished owing to the hydrophobic nature of the formulation.
All of this translates to an extended working time for
accomplishing the manufacturing steps that are required to make a
composite delivery device, including the time needed to complete
multiple process steps, such as mixing, metering, impregnating,
conveying, cutting, and packaging.
[0141] On the other hand, the creation of a composite reinforced
delivery device including a pharmaceutical formulation comprising a
hydrophilic component poses a different set of challenges.
Importantly, from a process manufacture perspective, if fiber
reinforcement is to be employed, then it is preferable to intermix
and to pre-wet the cellulose fibrous components with the
pharmaceutical formulation prior to the onset of appreciable
gelation. This is because the fibers can be more easily wetted and
intermeshed with the formulation when the formulation exhibits low
viscosity and minimal elastic recovery as it would prior to
gelation. In order to accomplish this process step, there needs to
be ample working time prior to gelation to facilitate the total
time requirements for vehicle mixing, metering, wetting, and
infiltration or impregnation of the fibrous material. For example,
when water is mixed with GLBG at a 2/1 (w/w) ratio, gelation and
elastic network formation is observed to begin almost immediately.
However, for the case of glycerin, the work time window prior to
the onset of gelation was observed to be significantly longer,
thereby making glycerin a more practical choice as a liquid for
creating hemostatic fiber-reinforced hydrophilic devices.
[0142] It is understood by those skilled in the art that within
some time-period after mixing liquids like water or glycerin with
gelatin, gelation will begin to occur, and the initial suspension
of discrete gelatin particulates will become transformed into an
elastic gelled network of surface-bonded, aggregated gelatin
particulates. The time-period preceding gelation is herein referred
to as "the work-time" and defines the window of time that enables
the delivery device to be made through the process of impregnating
a fibrous substrate. As long as the process is initiated during the
work-time, prior to gelation, the viscosity and elasticity of the
formulation will be low enough to enable facile impregnation of
fibrous substrates with high expediency. Thus, it is desirable that
the gelation process be made to occur after the fibrous textile is
impregnated with the formulation, and not before.
[0143] For the purposes of creating a fiber-reinforced delivery
device with a pharmaceutical formulation comprising a hydrophilic
component, it is desirable that the liquid component be miscible
enough with the hygroscopic network-forming component to lead to
gelation and to the formation of a plasticized polymer network,
including gums like gelatin, gum arabic, ghatti, karaya,
tragacanth, agar, Irish moss, carrageenin, alginates, seed extracts
of which include locust bean, locust kernel, and quince seed gums
as examples, manufactured and modified dextrins and British gums,
water-dispersible or soluble derivatives of cellulose, etc. It is
further desirable that the work-time prior to gelation be long
enough to facilitate all of the process steps that are required for
product formation, such as vehicle mixing, metering, conveying,
wetting, or pressing. If a continuous or semi-continuous process is
used to meter and convey the formulation onto a web of fibrous
material, such as the cellulose hemostat, then the web could be
optionally conveyed through a forced air or infrared heated oven to
facilitate faster gelation. Regardless of the use of ovens, once
the gelation process is complete, the resulting vehicle-impregnated
composite can be cut to achieve the desired geometric size for the
application, and then the resulting delivery device can be packaged
for storage prior to deployment. If an additive manufacturing
process is used to meter and convey the formulation onto a web or
discrete sheets of fibrous material, such as the cellulose
hemostat, then the formulation could be propelled from a three
dimensional printer nozzle or printer jet onto the web or discrete
sheets, resulting in vehicle impregnated composites of the desired
geometric size for the application, and then the resulting delivery
device can be packaged for storage prior to deployment.
Three-dimensional printing would also result in the creation of
customized dose and dosage forms impregnated into the reinforcing
member if so desired.
[0144] Regarding storage, it is further desirable that the liquid
component be biostable, either on its own, or through the
incorporation of preservatives that guard against bacterial growth
during periods of product manufacturing, packaging and storage. It
is also desirable that the liquid lead to formation of a gelled
polymer network after textile impregnation and not before. One
example of a liquid that meets both criteria is glycerin. Other
liquids can be used, including for example, propylene glycol,
polyethylene glycols and polypropylene glycols of various molecular
weights, water-based natural products like honey, polyhydric
alcohols and derivatives of the same, as well as mixtures of any of
these types.
[0145] It is also important that the fibrous components of the
composite delivery device be resistant to deterioration, swelling,
or dissolution by the hydrophilic liquid. SURGICEL.RTM. Original
Absorbable Hemostat (SO) textiles were determined to be resistant
to glycerin. In a separate experiment, pre-cut SO textiles
((1.8.times.3.8 cm) were separately drop-coated with glycerin and
water. After 24 hours, the glycerin-coated textile was observed to
retain its meshed structure with no noticeable evidence of
dissolution or physical changes (e.g., no shrinkage or swelling).
In a similar test, the SO textile was also observed to be more
resistant to water than its SafeGauze counterpart. SafeGauze
dissolved upon exposure to water as shown in Example 5, whereas SO
showed no apparent signs of dissolution within a 24-hour window of
testing (only shrinkage).
[0146] Regardless of whether the pharmaceutical formulation
comprises hydrophobic or hydrophilic components, the resistance of
the fibrous material to water dissolution or to degradation can be
an important and desirable attribute, particularly after
deployment. Although it is desirable that the fibrous material
eventually degrade and become bio-absorbed, it is still desirable
that the fibrous material maintain integrity for a period of time
during the post-deployment lifetime, mainly because the retention
of a composite structure with fibrous reinforcement is conducive to
maximizing macroscopic erosion resistance, which is another
desirable attribute for longer-term durability if deployed in a
tooth extraction socket application.
[0147] In one embodiment of the sustained release pharmaceutical
formulation, a solid, flexible pain management delivery device
comprises a mixture of 10-20% of a network-forming binder material
and 80-90% bupivacaine-loaded PLGA microspheres, wherein the
mixture is impregnated within a fibrous matrix material, such as a
flexible water-soluble cellulose fiber textile. The network-forming
binder material may comprise one or more components alone or in
combination, including a wax component (e.g., carnauba, palm,
beeswax), a gelatin component (e.g., GLBG, GLPG, SF), and an oil
(e.g., mineral oil or soy or palm oil). The mixture may further
comprise an optional pH modulator (e.g., citric acid,
di-Na-citrate). If a wax is employed, it is preferred that it be
ingestible. Oils and extenders should be USP-grade and also
ingestible. PLGA average particle sizes can range between 1 and 80
microns, with active ingredients comprising 1 to 50% by weight of
the PLGA encapsulated particulates, and where one preferred PLGA
particle size distribution comprises maximizing the % of small
particles (e.g., 3.4 micron) while simultaneously balancing all of
the aforementioned considerations for controlling drug-release
profiles as in the sustained release formulation comprising the
powdered mixture embodiments.
[0148] Thus, in order to maximize solids while simultaneously
providing a network-forming material binder component (e.g., GLBG)
capable of binding PLGA spheres upon hydration, it is desirable to
maximize the particle size of the ground gelatin component. This
mixture comprises 83.72% total solids in mineral oil, capable of
delivering 206 mg bupivacaine to a 0.55 cc cavity.
[0149] Determining the optimum level of wax/oil required to
facilitate textile-impregnation requires consideration of 1)
compliance of the mixture, 2) cohesive strength of the mixture, 3)
in vivo hydration rate of the mixture, 3) mechanical properties of
the mixture as a function of time during the in vivo hydration
process, 4) conduciveness of the mixture to textile impregnation
processes (e.g., solvent-free, minimal pressure, minimal
temperature, textile wettability, etc.), and 5) capacity to
pre-hydrate with water before insertion into the tooth extraction
socket.
[0150] A pharmaceutical formulation is possible with a low melting
wax, or with an oil/wax blend, or with a low Tg polymer (lower than
the Tg of the PLGA). A simple pressing process may be used to
consolidate the textile with the PLGA spheres under ambient
conditions. Optionally, gelatin may be omitted from the formulation
to thereby allow the cellulose to become the binder when it
hydrates. Omission of the gelatin would make more "room" for
bupivacaine loaded PLGA microspheres.
[0151] Selective surface-active molecules or surfactants can be
optionally incorporated into the mixture for the purpose of further
controlling the batch-to-batch consistency and rheological
characteristics of the pharmaceutical formulation to the degree
necessary for achieving desired reproducibility, tactile feel, and
efficacy. Such additives can be used for stabilizing oil-in-water
dispersions or emulsions, water-in-oil dispersions or emulsions,
and solid-in-oil dispersions. Surface active agents with surfactant
properties can include additives such as lecithin, fatty acid
esters, non-ionic polymers (e.g., polyvinyl alcohol), and the like.
Optional surfactants can include those known to the art, including
those where the overall effective HLB value is conducive to the
formation of water-in-oil emulsions (HLB<6), and those conducive
to the formation of oil-in-water emulsions. The amount should be
about 0.15 to 5.0 weight percent of the composition, and preferably
0.5 to 2.0 percent by weight. As is well known to those of ordinary
skill in this art, the HLB value is determined by a standardized
technique for measuring the solubility of a surfactant. Said
surfactant may be anionic, cationic or non-ionic with respect to
its hydrophilic portion. However, it is preferable that the
surfactant be biocompatible and ingestible. Examples of surfactants
can include polysorbates, mono-fatty acid esters of polyoxyethylene
sorbitan such as Tween-20 and Tween-80, polyglycerol
polyricinoleate, monoglycerides, lecithins, citric acid esters,
glycolipids, fatty alcohols and fatty acids, ethoxylated
polyhydroxystearate esters, glyceryl monooleate, polyglyceryl
esters, sorbitan esters, and propylene glycol esters. Other
examples of ingestible surfactants known in the art can be found in
RK Sharma, Surfactants: Basics and Versatility in Food Industries,
PharmaTutor, 2014, 2(3), 17-29.
[0152] These types of additives can be optionally incorporated
within one or two stages: (1) during the PLGA particle
manufacturing process, or (2) separately, during the compounding
process within the formulated vehicle to the degree required to
facilitate efficient rheological control for compounding, for
subsequent optional textile-impregnation, or for rheological
properties after hydration. The decisions regarding these additives
will be primarily based on or weighted by rheological responses
associated with manufacturing, for example shear-dependent
viscosity, and on the tactile-feel of the product in its end use,
in particular viscosity prior to hydration and compliance after
hydration.
[0153] Statistical formulation design of experiments (DOE's) can be
used to make weighted use of the aforementioned factors to modulate
release profiles. The release profile responses can then be modeled
along with rheological responses to achieve a navigable design
space as a function of all formulation factors for the ultimate
co-optimization of the response sets, co-optimization of release
responses and rheological responses that impact manufacturability
and end use tactile characteristics. Standard operating procedures
for the compounding and manufacturing process will insure
achievement of a consistent state of dispersion within the
optimized formulated product. This consistency in raw materials and
manufacturing processes will be paramount to product consistency,
reliability, and efficacy.
[0154] For the case of pharmaceutical formulations employing
hydrophobic components, oils are can be premixed with a wax at
elevated temperatures above the melt temperature of the wax to form
solutions. The solutions are then allowed to cool, causing a
portion of the wax to recrystallize into micro-crystallites, which
then remain suspended within the oil carrier. The resulting
mixtures of oil and wax have higher viscosity than neat oil and are
therefore desirable for use in formulating stable dispersions of
particulates that can resist settling over time. It can be
appreciated that the rate of cooling can be used to modulate the
size of the resulting micro-crystallites, with fast cooling
generally leading to smaller crystallites, and with slow cooling or
annealing leading to larger crystallites. Either way, the purpose
is to yield gelatinous mixtures, which serve as vehicles for
suspending particulates of network-forming polymers and active
ingredients within the pharmaceutical formulation. The viscosities
of gelatinous mixtures of oil and wax may be modulated by changing
the ratio of oil to wax, the wax type, and the oil type. It is also
possible to mix combinations of different types of waxes with
different types of oils at different ratios. One of the advantages
of the latter approach can be to minimize the level of oil in the
gelatinous mixture and hence in the final formulation. Mixtures of
hydrophobic waxes and oils can be used as components of the binder
material in a formulation for impregnation into a cellulose
textile.
[0155] Several types of oils or mixtures of oils could also be used
in combination with a wax at ratios of total oil to wax that are
sufficient for enabling certain desirable physical attributes,
including melting point depression of the wax, increase in
compliance of the resulting mixture, compressibility of the
resulting mixture for textile impregnation, temperature-dependent
viscosity of the resulting mixture, % PLGA and % gelatin loading in
the mixture, and the like. The typical oil to wax w/w ratio can be
in the range of 0.01/1 (still solid and wax-like) to 10/1 (weak
gelatinous amalgam) or higher.
[0156] The choice of oil type and oil amount will also depend on
other types of physical-chemical factors. Examples of such factors
include: 1) the degree that it is desirable to minimize the total
oil level in the final formulation mixture; 2) the threshold level
of oil needed to impart a sense of flavor if desired; 3) the
threshold level of an oil needed to impart analgesic effects; and
3) the solubility characteristics of other desirable solid active
ingredients within the oil phase.
[0157] Examples of oil types that can be used alone or in
combination include mineral oil, isopropyl palmitate, caprylic
triglyceride, coconut oils, vegetable oils like soy oil, corn oil,
sunflower oil, castor oil, and canola oil, aloe, apricot, argan,
avocado, camelina, D-limonene, olive oils, grapeseed oil, hempseed
oil, palm oil, rice bran oil, rosehip oil, safflower oil, sesame
oil, soy lecithin, almond oil, tamanu oil, vitamin E, walnut oil,
wheat germ oil, fish oils, and others.
[0158] Examples of oils or infused oils that can be used alone or
in combination with oils mentioned above to impart flavor or
analgesic effects include, for example, oils of spearmint,
peppermint, wintergreen, clove, cinnamon, palo santo, lavender,
juniper, oregano, thyme, geranium, ginger, nutmeg, pine, rose,
nutmeg, clove, coriander, citronella, lemon, anise, tea tree,
orange, turmeric, allspice, ho wood, cypress, and eucalyptus as
reported by Silva, J. et al., in the Journal of Ethnopharmacology,
Volume 89, Issues 2-3, December 2003, Pages 277-283.
[0159] A simplified manufacturing process for the hydrophobic
embodiment would involve a continuous process method comprising the
steps of 1) a carrier (e.g., a release-lined paper) is coated with
any of the wax-based embodiments described herein (with or without
the addition of dispersed gelatin particles) to yield a coated
continuous web; 2) the tackiness of the wax-based coating is made
sufficient so as to facilitate contact adhesion with PLGA particles
(either by means of formula ratios, temperature, or both); 3)
microencapsulated particles containing active ingredients such as
BUPIVACAINE are metered and distributed uniformly along the moving
web, or the web is moved through a fluidized bed of
microencapsulated particles to achieve contact adhesion of the
particles to the web with optional self-minimization of deposition;
4) the knitted, woven or non-woven cellulose material is
contact-pressed over the moving web with press rollers; 5) the
resulting composite is optionally die-cut into prescribed shapes
and weights; 6) the release material is peeled away from the
device, or is allowed to remain intact before the device is
packaged. The release liner material could even be synonymous with
the lower member of the package for the device, where the upper
component member of the package could be another type of release
layer that is used to sandwich and form a seal around the device
during packaging under sterile conditions.
[0160] Another manufacturing process for the hydrophobic embodiment
would involve an additive manufacturing process method comprising
the steps of 1) a knitted, woven or non-woven cellulose textile is
sized, cut and secured to fit the printing bed of a
three-dimensional printer; 2) the formulation comprised of the
gelatin, the microencapsulated particles containing active
ingredients such as BUPIVACAINE, and a wax or oil vehicle are
propelled through a moving three-dimensional printer nozzle or jet
to distribute uniformly along the stationary textile; 3) different
ratios of active ingredients, gelatin and hydrophobic additives
could be altered in a programmed fashion to produce a variety and
customized range of active ingredient doses and dosage forms across
a single sheet; 4) the resulting composite is optionally die-cut
into prescribed shapes and weights. Additionally, the reinforcing
member could be in the form of a particle or chopped fiber, added
directly to the formulation mixture and three-dimensionally printed
in a similar fashion into discrete sheets to be packaged and
sterilized for use.
[0161] Certain hydrophobic formulations have been observed to have
rheological characteristics that are conducive to the use of a
sigma-blade blending process for preparing mixtures under ambient
conditions, whereby the PLGA powders and gelatin particulates could
be added to form dough-like mixtures in a batch or semi-continuous
batch process. For formulations comprising mixtures of wax and oil,
melt-recrystallization steps could also be optionally employed to
produce stiffer or less stiff mixtures upon cooling. In addition,
shear mixing of hydrophobic formulas could be performed with the
intent of generating shear-induced heat. The resulting process
temperature could be controlled and maintained at temperatures of
less than the Tg of PLGA, and the mixture could then be directly
dispensed onto a textile for impregnation while cooling. Dispensing
could be done directly into a kit comprising a blister package
containing pre-inserted and precut textiles, or onto a larger
textile, which would then be subsequently cut into desired
dimensions for end use. A compression step could be employed to
help insure impregnation of the interstitial spaces if
necessary.
[0162] In one embodiment of the above device, a solid, flexible
formulated pain management device comprises a mixture of 10-20% of
a binder material and 80-90% bupivacaine-loaded PLGA microspheres,
wherein the mixture is impregnated within a fibrous matrix
material, such as a flexible water-soluble cellulose fiber textile,
wherein the binder material is comprised of one or more components
alone or in combination. The binder material may comprise a wax
component (e.g., carnauba, palm, beeswax), a gelatin component
(e.g., GLBG, GLPG, SF), and an oil (e.g., mineral oil, soy oil, or
palm oil as optional hydrophobic component). The mixture may
further comprise an optional pH modulator (e.g., citric acid,
di-Na-citrate). If a wax is employed, it is preferred that it be
edible/ingestible. Oils/extenders should be USP-grade and also
ingestible. A preferred PLGA particle size distribution comprises
maximizing the % of small particles (e.g., D50=3.4 micron) while
simultaneously balancing all of the aforementioned considerations
for controlling drug-release profiles as mentioned previously in
discussions pertaining to the powdered mixture embodiments of
hydrophilic devices. In Examples 3 through 8, a 30/70 w/w blend of
3.4 um and 42.7 um PLGA microspheres were used to demonstrate the
concepts leading to the formulation of a unique hydrophobic
controlled-release delivery device.
[0163] An embodiment of a manufacturing process for a
pharmaceutical formulation comprising a hydrophobic component
involves the use of a continuous web coating process method whereby
a moving carrier, such as a wax or silicone release-lined paper, a
knitted, woven or non-woven hemostatic textile, etc., is first
coated with any of the wax-based mixtures as described herein,
optionally with or without the addition of dispersed gelatin
particles to yield a coated continuous web. The tackiness of the
wax-based coating on the moving web is intended to facilitate
contact-adhesion with PLGA particles containing bupivacaine (BUP).
The PLGA particles are metered and distributed uniformly along the
moving web using mechanisms such as passing the web through a
fluidized air chamber, drop-metering PLGA powder directly onto a
moving web, spinning-disc dry-metering PLGA particles as they are
being formed during the microencapsulation manufacturing process,
or passing the web through a fluidized bed of PLGA particles to
achieve contact adhesion between the particles and the web with
optional self-minimization of deposition. A knitted, woven or
non-woven cellulose hemostat material is then optionally
contact-pressed over the moving web with press rollers. The
resulting composite is optionally die-cut into prescribed shapes
and weights. If the moving carrier is a wax or silicone coated
release paper, then the paper is optionally peeled away from the
device, or it is allowed to remain intact before the device is
packaged, or if the moving carrier is a fibrous hemostat, then the
entire impregnated assembly is packaged.
[0164] In another embodiment of the manufacturing process, the
moving carrier could be an integral component of the product
package itself. In this case, a heat-sealable material like
polyethylene terephthalate (PET), high density polyethylene (HDPE),
or a foil laminate could be used as the first moving carrier, which
is then coated with the wax-based amalgams as described herein, and
then passed through any one of the PLGA coating processes as
described above. The carrier is then contact pressed with a
knitted, woven or non-woven hemostat textile, followed by contact
pressing with an upper component package layer such as PET, a foil
laminate, or an HDPE film. The composite is then finally die-cut
and pressure-sealed with optional heat to yield the final packaged
device.
[0165] Hydrophobic formulations containing amalgamized dispersions
of particulates in mixtures of oil and wax, such as gelatin
particulates, particulates of active ingredients, and particulates
of active ingredients encapsulated with encapsulating materials,
exhibit rheological characteristics that are conducive to the use
of batch mixing processes under ambient conditions (e.g., a
sigma-blade blending process). In one process, mixtures of oil and
wax are first prepared, and then amalgams are prepared by blending
the pre-mixed oil/wax mixtures with gelatin particulates, and then
the PLGA powders are metered into the amalgams to form dough-like
mixtures. Similarly, continuous mixing processes could also be
employed, such as single screw or twin-screw extruders with
appropriate mixing zones and metering zones for continuous shear
mixing under near-ambient conditions, followed by an exit die for
cutting and metering aliquots of the mixtures onto a continuous
moving web for packaging. A melt-recrystallization step could also
be optionally employed, which would likely lead to stiffer mixtures
upon cooling. In addition, shear mixing of higher viscosity
formulations could be performed with the intent of generating
shear-induced heat if so desired. The processing temperature could
then be controlled and maintained with air or liquid cooling, so
that the temperature of the mixture remains lower than the glass
transition of PLGA (e.g., the Tg of RG504 PLGA is 46-50 degrees C.)
in order to minimize premature process-induced BUP diffusion. The
mixture could then be directly dispensed onto a textile while
molten and hot for easier impregnation while cooling. The delivery
device could then be directly dispensed into a kit comprising a
blister package with pre-inserted individual textiles, or directly
onto a larger continuous moving web of textile, which could then be
subsequently cut into desired dimensions for end use. A compression
step could be employed to help insure impregnation of the
interstitial spaces if necessary. For formulations that are
deployed without fibrous reinforcement, metering and dispensing of
the formulations could be performed directly into blister packages
for subsequent sealing, shipping, and storage.
[0166] The manufacturing processes as described above for
pharmaceutical formulations comprising hydrophobic ingredients
would not be applicable to pharmaceutical formulations comprising
hydrophilic ingredients of the types described in Examples 1 and 2.
However, for pharmaceutical formulations comprising hydrophilic
ingredients of the types described in Example 15, any of the
aforementioned batch or continuous processes could be similarly
adapted and employed for the manufacture of hydrophilic delivery
devices.
[0167] As demonstrated in Example 1, certain additives like citric
acid can have a positive impact on the gelation characteristics and
on the mechanical property characteristics of binder components.
These same additives can also be used to modulate the chemical
environments within hydrophobic and hydrophilic devices;
particularly during in vivo hydration, where fluids such as saliva
and blood can diffuse from the tooth extraction socket into the
device, and active ingredients can diffuse out of the device.
[0168] The overall impact of pH modulators can be to alter
diffusion characteristics via at least two different mechanisms: 1)
by impacting the solubility of active ingredients within the
device; and 2) by altering the mechanical properties of the
diffusing medium which in turn impacts free volume and the
resulting rate of molecular-scale diffusion through the medium.
[0169] The effects of pH on gelatin binder properties can be
appreciated with the results presented in Example 1, where unlike
formulas made under pH-neutral conditions, the relative viscosities
of formulas made with citric acid were observed to undergo
significant changes within five hours of mixing, and even more so
within one day after mixing.
[0170] Certain gelatins were observed to form low-modulus elastic
networks (plasticized with water) at faster rates in the presence
of citric acid than in the absence of citric acid. It follows that
by controlling the ratio of citric acid (or other alternative
acids) to protein, the rate of network gelation can be modulated,
implying that the mechanical resistance to diffusion can be
similarly modulated and controlled. Higher citric to gelatin ratios
would likely lead to even faster gelation rates (as would lower
levels of water).
[0171] Similarly, a hydrophobic device can also be formulated with
a pH modulator such as citric acid or di-sodium citrate. As
demonstrated in Example 13, this can be accomplished by dispersing
the ingredients as powders within the formulation vehicle. Given
that hydrophilic additives will have limited solubility in the oil
carriers that are used in hydrophobic devices, the additives will
remain dispersed and undissolved until the device becomes hydrated,
either in vivo, or prior to deployment. However, it is also
possible to incorporate oil-soluble or partially oil-soluble protic
acids into hydrophobic devices if so desired (e.g., long-chain
fatty acids such as stearic acid, lauric acid, sebacic acid,
etc.).
[0172] Importantly, any of these types of pH-modulator additives
can be optionally microencapsulated with polymers like PLGA, or
with other types of polymers, to control solubility and rate of
release within the end use chemical environment.
[0173] Simultaneously, it can be appreciated that a balance would
have to be achieved with other factors that impact the efficacy of
the device, including the aforementioned chemical environment
factors that affect the solubility and diffusion of bupivacaine
(e.g., the initial concentration of bupivacaine in its free-base
and in its acidic form; and the ultimate equilibrium concentration
of both species within the end use chemical environment).
[0174] Hydrophobic oils and mixtures of oils with waxes can be
considered as carriers for various dispersed ingredients that
constitute the binder phase of a delivery vehicle for a hydrophobic
formulation or delivery system. Again, examples of dispersed
ingredients can include one or more of gelatin particulates, other
network forming polymers, PLGA microspheres containing active
ingredients like BUP free base or BUP-HCl, other types of active
ingredients, including those encapsulated with PLGA, those
encapsulated with alternative encapsulating materials, or those
with no encapsulant, citric acid powder, di-sodium citrate powder,
BUP free-base, BUP-HCl, and others. Many types of oils, oil
mixtures, or oil/wax mixtures can be employed, provided that they
meet the criteria for use in a hydrophobic device as described by
tests presented in Example 14. Examples of oils that satisfy these
criteria include mineral oil as described in Example 8, isopropyl
palmitate or caprylic triglyceride as described in Example 10, or
coconut oil as described in Example 16.
[0175] Again, these types of oils or others can also be used in
combination with waxes to modify the rheological characteristics of
the liquid carrier, and to modify the rheological and mechanical
compliance characteristics of the resulting device. Examples of
waxes include carnauba wax, beeswax, paraffin wax, palm wax and
mixtures thereof as described in Example 8, as well as others. A
wax or wax-mixture is typically employed at levels such that the
total oil to wax ratio facilitates the achievement and control of
certain desirable physical attributes or property characteristics,
including, for example: 1) melting point depression of the wax; 2)
an increase or decrease in the compliance of the resulting vehicle;
3) an increase in the cohesive integrity of the vehicle; 4) an
increase or decrease in the viscosity of the resulting vehicle,
such that dispersed ingredients within the vehicle remain dispersed
without settling; 5) achievement of vehicle compliance with minimal
elastic recovery to facilitate textile-impregnation processes; 6)
achievement of a temperature-dependent viscosity characteristics
that are conducive to shear mixing processes for vehicle
manufacturing; 7) achievement of the ability to control or maximize
the % PLGA and hence the % BUP in the vehicle, which can increase
the dosage potential of the vehicle if so desired; and 8)
achievement of the ability to control or maximize the % gelatin
loading in the vehicle if so desired. The typical oil to wax w/w
ratio for satisfying these purposes can be in the range of 0.01/1
(still solid wax-like) to 10/1 (weak gelatinous amalgam) or even
higher if so desired.
Example 1. Limitations Pertaining to the Preparation of Hydrophilic
Devices
Part-A: Evaluation of Protein Binders
[0176] A series of commercially available materials were evaluated
for their relative water solubility, gelation, and network-forming
characteristics, including: 1) bovine collagen powder available
from Great Lakes Gelatin Company, Grayslake, Ill., Kosher, 100%
hydrolyzed collagen hydrolysate from bovine hide, >90% protein,
bloom 0 g, viscosity 5.5-7.5 mPa-s, pH 5.0-6.5, US Pharmacopeia
consumer grade; 2) piscine collagen powder available from Zint LLC,
(referred to as Marine Collagen, type-1 hydrolyzed fish collagen);
3) bovine gelatin (GLBG) powder (Great Lakes Gelatin Company,
Grayslake, Ill., type B, unflavored Kosher beef hide, bloom 225 g,
viscosity 34-40 mp, pH 4.1-5.5, 88-92% protein, US Pharmacopeia
consumer grade; 90% mesh estimated to be between about 35 and 70
(i.e., 0.5 mm to 0.2 mm)); 3) porcine gelatin (GLPG) powder
available from Great Lakes Gelatin Company, Grayslake, Ill., type
A, unflavored, 88-92% protein, bloom 225 g, viscosity 34-40 mp, pH
4.3-5.7, US Pharmacopeia consumer grade; and 4) Surgifoam (SF)
absorbable gelatin powder made from absorbable porcine gelatin
sponge, U.S.P., available from Ferrosan Medical Devices,
distributed by Ethicon, Inc.
[0177] Using either a spatula, a vortex mixer, or both, each of the
gelatin powders was hand-mixed with distilled water under ambient
conditions in separate glass containers at selected weight ratios
(w/w water to powder): 1/1, 2/1, 3/1, 10/1, 15/1, and 25/1. The
samples were qualitatively compared at selected time periods (t)
after mixing: 5 minutes, 15 minutes, 1 hour, 5-6 hours, 24 hours, 4
days, and 2 weeks for their relative viscosity characteristics as
gauged by their pourability and by their resistance to stirring
with a spatula. Collagen samples were similarly mixed and
evaluated, but the weight ratios of water to powder were limited to
1/1, 2/1, and 3/1 (w/w) owing to their poor network-forming
characteristics and lack of gelation. SF samples were evaluated at
the 3/1 and 25/1 w/w water to powder ratios for comparison to the
two other gelatin types from Great Lakes (i.e., GLPG and GLBG). pH
values were measured for each of the 25/1 (w/w) water/gelatin
samples at t=24 hours after mixing.
[0178] In general, the relative viscosity characteristics of all
samples were observed to increase with decreasing levels of water.
Several days after mixing the samples, the commercial collagen
samples were observed to exist as either low viscosity liquid
solutions or as liquid dispersions and remained pourable. Each of
the collagen samples remained pourable even at t=2 weeks after
mixing. Qualitative observations pertaining to the collagen samples
are provided in Table 1-1.
TABLE-US-00001 TABLE 1-1 Relative viscosity characteristics of
collagen samples after mixing with distilled water under ambient
conditions at weight ratios of water to collagen (w/w) of 1/1, 2/1,
3/1, 10/1, 15/1, and 25/1. The samples were qualitatively compared
at various time periods after mixing, including 5 minutes, 6 hours,
24 hours, 4 days, and 2 weeks. Sample Type and w/w water to 5
minutes 6 hours 24 hours 4 days after 2 weeks collagen ratio after
mixing after mixing after mixing mixing after mixing Great Lakes
Bovine Pourable liquid, Pourable liquid, Pourable liquid, Pourable
liquid, Pourable liquid, Collagen 1/1 clear solution clear solution
clear solution clear solution clear solution Great Lakes Bovine
PPourable liquid, Pourable liquid, Pourable liquid, Pourable
liquid, Pourable liquid, Collagen 2/1 clear solution clear solution
clear solution clear solution clear solution Great Lakes Bovine
Pourable liquid, Pourable liquid, Pourable liquid, Pourable liquid,
Pourable liquid, Collagen 3/1 clear solution clear solution clear
solution clear solution clear solution Zint Collagen 1/1 Pourable
liquid, Pourable liquid, Pourable liquid, Pourable liquid, Pourable
liquid, hazy dispersion, hazy dispersion, hazy dispersion, hazy
dispersion, hazy dispersion, partial solution partial solution
partial solution partial solution partial solution Zint Collagen
2/1 Pourable liquid, Pourable liquid, Pourable liquid, Pourable
liquid, Pourable liquid, hazy dispersion, hazy dispersion, hazy
dispersion, hazy dispersion, hazy dispersion, partial solution
partial solution partial solution partial solution partial solution
Zint Collagen 3/1 Pourable liquid, Pourable liquid, Pourable
liquid, Pourable liquid, Pourable liquid, hazy dispersion, hazy
dispersion, hazy dispersion, hazy dispersion, hazy dispersion,
partial solution partial solution partial solution partial solution
partial solution
[0179] By contrast, the gelatin samples exhibited significantly
higher relative viscosities than their collagen counterparts at
equivalent water to solid ratios. Unlike the collagen samples, the
gelatin samples were also observed to form high viscosity gelled
networks, where the time to gelation was generally observed to
increase with increasing levels of water.
[0180] Importantly, gelatin protein polymers with these types of
network-forming characteristics are preferred over their collagen
counterparts for use in the pharmaceutical formulations as
described herein. These types of network-forming polymers exhibit
mechanical property and cohesive strength characteristics that
render them as acceptable for use as binder-phase components for
other dispersed ingredients to be discussed in subsequent examples.
In this example, the gelatin proteins with Bloom values like those
reported for GLBG and GLPG (e.g., 225 g) are preferred over their
counterparts with lower Bloom values, such as the collagen
proteins. Qualitative observations pertaining to the gelatin
samples are provided in Table 1-2.
TABLE-US-00002 TABLE 1-2 Relative viscosity characteristics of
bovine and porcine gelatin samples after mixing with distilled
water under ambient conditions at weight ratios of distilled water
to gelatin (w/w) of 1/1, 2/1, 3/1, 10/1, 15/1, and 25/1. The
samples were qualitatively compared at select time periods after
mixing (t), including 5 minutes, 15 minutes, 1 hour, 6 hours, 24
hours, 4 days, and 2 weeks. Note that pH measurements were taken
for the 25/1 samples at t = 24 hours after mixing. Sample Type and
w/w water 5 minutes 15 minutes 1 hour 6 hours 24 hours 4 days 2
weeks to gelatin ratio after mixing after mixing after mixing after
mixing after mixing after mixing after mixing Great Lakes Bovine
Gelation, not Increasing Increasing Elastic gel No change No change
Gelatin 1/1 pourable gel stiffness gel stiffness network Great
Lakes Bovine Gelation, not Increasing Increasing Elastic gel No
change No change Gelatin 2/1 pourable gel stiffness gel stiffness
network Great Lakes Bovine Gelation, not Increasing Increasing
Elastic gel No change No change Gelatin 3/1 pourable gel stiffness
gel stiffness network Great Lakes Bovine Gelation, not Increasing
Increasing Increasing Elastic gel No change Gelatin 10/1 pourable
gel stiffness gel stiffness gel stiffness network Great Lakes
Bovine Gelation, not Increasing Increasing Increasing No change No
change Gelatin 15/1 pourable gel stiffness gel stiffness gel
stiffness Great Lakes Bovine Hazy dispersion, Hazy dispersion, Hazy
dispersion, Hazy dispersion, Hazy dispersion, Weak gel, Gelatin
25/1 pourable liquid pourable liquid pourable liquid pourable
liquid, pourable liquid Pourable pH = 5.8 after shaking Great Lakes
Porcine Gelation, not Increasing Increasing Elastic gel No change
No change Gelatin 1/1 pourable gel stiffness gel stiffness network
Great Lakes Porcine Gelation, not Increasing Increasing Elastic gel
No change No change Gelatin 2/1 pourable gel stiffness gel
stiffness network Great Lakes Porcine Gelation, not Increasing
Increasing Elastic gel No Change No change Gelatin 3/1 pourable gel
stiffness gel stiffness network Great Lakes Porcine Gelation, not
Increasing Increasing Increasing No change No change Gelatin 10/1
pourable gel stiffness gel stiffness gel stiffness Great Lakes
Porcine Hazy dispersion, Hazy dispersion, Hazy dispersion,
Gelation, No change No change Gelatin 15/1 pourable liquid pourable
liquid pourable liquid not pourable Great Lakes Porcine Hazy
dispersion, Hazy dispersion, Hazy dispersion, Hazy dispersion, Hazy
dispersion, Weak gel, Gelatin 25/1 pourable liquid pourable liquid
pourable liquid pourable liquid, pourable liquid Pourable pH = 5.19
after shaking Surgifoam 3/1 Gelation, not No change Non-elastic
Highly pourable compliant gel compliant gel Surgifoam 25/1 High
compliance, No change No change, No change No change high viscosity
pH = 5.25 gel, moves with shaking but does not pour
[0181] The Great Lakes porcine gelatin (GLPG) was observed to
produce lower relative viscosity mixtures than the Great Lakes
bovine gelatin (GLBG) at the same water/powder weight ratios. The
gelatin viscosity trends were similarly manifest at all
water/powder weight ratios up to a 25/1. After initial mixing, each
of the 25 to 1 water/gelatin samples were low-viscosity pourable
liquids. After several days, the 25/1 GLBG sample had become a more
homogeneous gel, but it was still pourable after shaking. The 25/1
GLPG sample was also still pourable, and it was lower in viscosity
than the 25/1 GLBG sample. The GLPG sample had also phase separated
into a partial gel with a clear supernatant. The two gelatin
protein types exhibited slightly different pH values from one
another at the 25 to 1 water/powder weight ratio.
[0182] In the next step, comparisons were made between Surgifoam
(SF) gelatin and the porcine and bovine gelatins using a 25/1
weight ratio of distilled pH-neutral water to powder and a 3/1
weight ratio of distilled pH-neutral water to powder. Unlike the
25/1 w/w GLBG and GLPG samples, the comparative 25/1 w/w SF sample
was observed to form a high viscosity non-pourable gel within 1
hour after mixing. By contrast, the GREAT LAKES bovine and porcine
gelatins formed hazy, lower viscosity dispersions, and they
remained pourable throughout a 4-day observation period. The 25/1
w/w SF sample was observed to have a slightly lower pH than the
comparative GLBG and GLPG samples.
[0183] When comparing the 3/1 w/w samples, the SF gelatin formed an
immediate gel which was highly compliant, pliable, and moldable. By
contrast, the GLBG and GLPC gelatins were much slower to gel than
SF. These trends were parallel to those observed when comparing the
25/1 w/w samples. After 1 day, the SF sample had become akin to a
dry-blend with some cohesive strength, and with the capacity for
much higher liquid adsorption. The 3/1 w/w SF sample was easily
broken with a spatula, and it was still high in compliance. By
contrast, the 3/1 w/w GLBG sample had become a fully cohesive
rubbery network with better cohesive strength and better cohesive
integrity than the 3/1 w/w SF sample. The GLPG sample behaved
similarly to the GLBG sample. In mechanical terms, the 3/1 w/w SF
sample retained a high degree of malleability and re-formability
with high compliance and low elasticity, whereas the GLBG and GLPG
samples exhibited higher elastic storage modulus characteristics.
After 4 days, these trends remained the same, SF was high in
compliance, and the GLBG and GLPG samples exhibited elastic
recovery. Thus, although the GLBG and GLPG samples were slower to
gel than SF, once they did gel, the GLPC and GLBG samples were more
elastic and less compliant than the comparative SF sample.
Part-B: Evaluation of Protein Binders with a pH Modulator (Citric
Acid)
[0184] The differences in pH among the three gelatin samples, SF,
GLBG, and GLPG, justified a separate test with citric acid to see
if the protein types would exhibit different degrees of acid
neutralization capacity, different relative rates of gelation, or
both. Using the same procedures as outlined above, a comparison was
made between the three gelatin samples at a 25/1 weight ratio of
distilled water to solids using a 1% citric acid solution having a
pH of 2.2. Unlike pH-neutral water, the slightly acidic citric acid
solution caused immediate partial-gelation of the proteins,
resulting in higher immediate relative viscosities for all three
samples, including Surgifoam.
[0185] The relative viscosity trends at t=1 hour after mixing were
the same as those observed under pH neutral conditions, but all
samples were higher in viscosity than those made without citric
acid. Qualitative observations at t=1 hour were recorded as
follows: SF was a cohesively weak gel that was moveable with
shaking but was not pourable; GLBG was a hazy gelled dispersion,
but it was still pourable; and GLPG was a hazy gelled dispersion
but it was still pourable. Qualitative relative viscosity trends
were recorded as follows: SF was much more viscous than GLBG which
was more viscous than GLPG.
[0186] The pH of the samples at t=1 hour after mixing were lower
than those of their counterparts that were mixed under pH-neutral
conditions, but slightly higher than 1% citric acid solution
itself, thereby providing evidence for some degree of acid
buffering and neutralization capacity. Thus, acid neutralization
via protein-amine protonation was observed to accompany the faster
rate of viscosity rise. The pH values for samples mixed with the 1%
citric acid solution were observed to trend similarly to those that
were measured under pH neutral conditions at t=1 hour. The
resultant pH values in the presence of 1% citric acid solution were
2.84 for the 25/1 w/w SF sample, 3.12 for the 25/1 w/w GLBG sample,
and 2.94 for the 25/1 w/w GLPG sample. The respective pH values for
the samples measured under pH neutral conditions were 5.25 for the
25/1 w/w SF sample, 5.80 for the 25/1 w/w GLBG sample, and 5.19 for
the 25/1 w/w GLPG sample.
[0187] At a time of t=5 hours after mixing, the relative viscosity
characteristics of the samples were observed to increase. The GLBG
sample was shakable and still pourable. The GLPG sample had turned
into a transparent gel and was shakable and still pourable. Thus,
protein-moiety protonation induced faster gelation and higher
relative viscosities in all three cases. Importantly, after t=24
hours, the trends were observed to become magnified. Although the
SF sample had remained unchanged as a high viscosity compliant gel
that was not pourable, the 25/1 w/w GLBG and GLPG samples had
become completely gelled. They were no longer shakable, nor were
they pourable. When a spatula was placed into the gels, the
relative viscosity trends were as follows: the GLBG sample was more
viscous than the GLPG sample which was only slightly more viscous
than the SF sample. By contrast, the comparative samples that were
mixed under pH-neutral conditions were observed to remain as
pourable liquids at t=24 hours, t=4 days, and t=2 weeks. Thus,
citric acid had not only successfully induced a faster rate of
gelation, it had facilitated a change in the relative viscosity
characteristics of the resulting gelled networks.
[0188] These results show that the acidity of the chemical
environment can be used to modulate the mechanical behavior of the
binder phase in formulations with gelatin proteins. As such, this
example demonstrates that acids, such as citric acid and others,
can be used as optional components into formulations for the
purpose of modulating gelation rates and mechanical property
characteristics of the resulting formulations. Importantly, this
result demonstrates that the pH of the chemical environment will
have an impact on the rheological characteristics of the
formulation. This in turn will not only have an impact on the
diffusion rate of active molecules like bupivacaine, but it will
also have an impact on the compliance characteristics of the
formulation, which in turn will affect its formability, or
compliance, when affixed within a static volume cavity such as a
tooth extraction socket.
Part-C: Statistically Designed Experiments (DOE) for Formulations
to Deliver Targeted Dosages of Bupivacaine (BUP)
[0189] A statistical DOE was performed for the purpose of
investigating the viability of providing a system to deliver a
targeted theoretical-maximum level of bupivacaine (BUP) into a
fixed volume cavity. A Taguchi 4-factor 3-level statistical design
template was chosen for this work owing to its ability to provide
maximum learning potential via trend analyses while simultaneously
preserving economy of scale, materials, and time. Multiple DOE
drafts were conceptualized with one of the overall objectives being
to use the data to develop a qualitative and cursory understanding
of the impact of poly(lactic-co-glycolic acid) (PLGA) particle size
distribution and gelatin binder content on the relative hydration
behavior, rheology characteristics, and compliance characteristics
of resulting devices. Observations were made in three stages: (i)
immediately after mixing, (ii) as a function of time after mixing,
and (iii) with additional hydration after mixing. Initial work was
done with placebo microspheres only. Note that 20% w/w bupivacaine
loaded PLGA microspheres were used in subsequent experiments for
analytical studies.
[0190] Initially, the target dosage range for bupivacaine was
estimated to be between a level approaching possible toxicity on a
high-delivery side and a level representing clinical usefulness on
a low-delivery side. The upper limit of BUP was estimated to be 360
mg over a 4-day period (90 mg/day.times.4). Importantly, because of
the unique fixed volume constraint in the end use application, the
theoretical formulation composition for the upper limit and the
lower limit for the % gelatin binder and the % of dispersed PLGA
microspheres were observed to be dictated by the bupivacaine target
dosage level. Due to the unique occupied volume limitation for this
type of end use application, estimated as 1 cc in this example, any
change in the weight % of the gelatin binder necessitates an
opposite change in the weight % of the PLGA microspheres. Given
that the PLGA microspheres function as the carriers for the active
BUP molecules, it follows that higher BUP dosages necessitate
higher percentages of PLGA microspheres and lower percentages of
binder.
[0191] Initially, attempts were made to achieve the theoretical
upper limit dosage of BUP by using 20% w/w loading of the PLGA
microspheres. However, with a maximum bupivacaine target of 360 mg
over a 4-day period (90 mg/day.times.4), the target was determined
to be not viable. In order to achieve the targeted upper dosage
limit with 20% w/w BUP-loaded PLGA microspheres in a fixed volume
cavity, the formulation would have to be prepared with little to no
binder. In the absence of binder, the device would have no cohesive
integrity, and the PLGA microspheres would easily erode away and
evacuate the tooth extraction socket. For this reason, calculations
were performed in an attempt to satisfy a condition whereby the
upper limit for BUP dosage would be delivered into a single tooth
extraction socket cavity estimated to be 1 cc in volume via a
formulation comprising PLGA microspheres dispersed within a
hydrated gelatin binder. The target was deemed to be achievable
only by increasing the theoretical % w/w BUP loading of the PLGA
microspheres to a level that enabled the use of a binder together
with additional of fluids for hydration of the device.
[0192] Three different pathways were identified to approach the
problem of maximizing dosage: (1) increasing bupivacaine-loading to
its maximum theoretical level of about 50% w/w within the PLGA
microspheres; (2) minimizing the gelatin binder levels to the
extent permitted without simultaneously deteriorating mechanical
properties; and (3) minimizing the level of water required for
hydration/mastication to the extent tolerable without experiencing
unmanageable decreases in compliance. With these limitations in
mind, "DRAFT-1 DOE" was created to study the effects of four
factors: [0193] FACTOR-1=weight percent solids in the hydrated
formula (PLGA microspheres+SF=50%, 53%, and 56%); [0194]
FACTOR-2=interchangeable choice between binder-type (SF, GLBG,
GLPG) or pH-modulator (standard pH water, 1% citric acid solution,
1% di-sodium citrate solution); [0195] FACTOR-3=the weight fraction
of small (D50=3.4 micron) PLGA particles as a percentage of all
PLGA particles (0, 0.05, 0.1); [0196] FACTOR-4=the weight fraction
of D50=42.7 micron spinning-disc-dried PLGA particles (0, 0.15,
0.3), wherein D50=42.1 micron emulsion PLGA particles constituted
the balance (e.g., 1, 0.8, 0.6).
[0197] Based on the DRAFT-1 DOE constraints, the maximum viable
dosage target for bupivacaine was theoretically determined to be
300 mg in a 1 cc fixed volume cavity, which was a level that was
closer to the original 360 mg dosage target. Although higher
bupivacaine dosages would theoretically still be possible, it was
recognized that an appropriate level of gelatin binder would still
be needed to hold the formula together during hydration. Again,
although PLGA microspheres were used for this experiment, it was
assumed that 50% bupivacaine-loading of the PLGA microspheres would
also be possible. Surgifoam (SF) powder was initially used as the
binder. Table 1-3 provides information on the PLGA microspheres
provided by Southwest Research Institute (SWRI). Tables 1-4 and 1-5
reveal pertinent DRAFT-1 DOE calculations based on the initial
constraints as described above. The gelatin binder and PLGA
microsphere powders were dry-mixed at the specified weight ratios,
and selected dry mixtures were then mixed with water at specified
weight ratios using a hand-held spatula.
TABLE-US-00003 TABLE 1-3 PLGA microsphere information from
Southwest Research Institute (SWRI). Bupivacaine Free Base Sample
ID Loading Amount of (NB: 18- (Theoretical Polymer Sample D (0.1) D
(0.5) D (0.9) 0202-015-) Wt %) Matrix Process/Comments (grams)
microns microns microns 5 0% Resomer Spray drying using 3.7 g 19.2
42.7 88.5 Placebo RG 504 spinning disk Low recovery due to
agglomeration and sticking inside of atomization chamber 6 20%
Resomer Spray drying using 0.6 g 24.5 52.1 101.4 RG 504 spinning
disk Low recovery due to agglomeration and sticking inside of
atomization chamber 7 0% Resomer Spray drying using two- 4.0 g 1.4
3.4 9.2 Placebo RG 504 fluid nozzle 10 20% Resomer Spray drying
using two- 5.0 g 1.0 3.5 7.0 RG 504 fluid nozzle 9 0% Resomer
Emulsion, Solvent- 4.4 g 27.8 42.1 63.4 Placebo RG504 extraction;
photo provided in FIG. 7f
TABLE-US-00004 TABLE 1-4 DOE DRAFT-1 specifications calculations
for creation of dry compositions and for calculations presented in
Table 1-5. FACTOR-1 FACTOR-3 FACTOR-4 Wt. % Solids Optional
Factor-2 weight fraction weight fraction weight fraction in
hydrated FACTOR-2 Hydration Solution of D50 3.4 um of D50 42.7 um
of D50 42.7 um Expt. formula Gelatin Type Type PLGA microspheres
PLGA microspheres PLGA microspheres 1 50.47% SF pH-neutral 0 0 1 2
50.47% GLBG 1% citric in water 0.05 0.15 0.8 3 50.47% GLPG 1%
Na-citrate in 0.1 0.3 0.6 water 4 53.47% SF standard 0.05 0.3 0.65
5 53.47% GLBG 1% citric in water 0.1 0 0.9 6 53.47% GLPG 1%
Na-citrate in 0 0.15 0.85 water 7 56.47% SF pH-neutral 0.1 0.15
0.75 8 56.47% GLBG 1% citric in water 0 0.3 0.7 9 56.47% GLPG 1%
Na-citrate in 0.05 0 0.95 water 10 53.47% SF pH-neutral 1 0 0 11
53.47% SF pH-neutral 0 1 0 12 53.47% SF pH-neutral 0 0 1 13 25.00%
SF pH-neutral 0 0 1
TABLE-US-00005 TABLE 1-5 DOE DRAFT-1 calculations of pertinent
composition information based on the constraints presented in Table
1-4. Although placebo PLGA microspheres were used in preparing
samples, calculations were performed to estimate a theoretical
dosage of BUP delivery to a tooth extraction socket of 1 cc volume,
assuming that the PLGA microspheres were loaded with 50% BUP by
weight. Formulation 1 2 3 4 5 6 Target Bupivacaine Dose over 4-Day
0.300 0.300 0.300 0.300 0.300 0.300 Period (grams) Estimated Tooth
extraction socket 1.3 1.3 1.3 1.3 1.3 1.3 Volume (cm.sup.3)
Estimated density of hydrated 1.1 1.1 1.1 1.1 1.1 1.1 formula
(g/cc) Estimated grams of hydrated 1.43 1.43 1.43 1.43 1.43 1.43
formula delivered to tooth extraction socket (g) Theoretical wt. %
Bupivacaine in 50% 50% 50% 50% 50% 50% microspheres Estimated
weight of drug-dosed 0.60 0.60 0.60 0.60 0.60 0.60 microspheres in
hydrated formula (g) Estimated % Total Solids in 50.47% 50.47%
50.47% 53.47% 53.47% 53.47% Hydrated Formula Wt. % Microspheres
dispersed in 41.96% 41.96% 41.96% 41.96% 41.96% 41.96% hydrated
formula Wt. % Gelatin in hydrated formula 8.51% 8.51% 8.51% 11.51%
11.51% 11.51% Wt. % Water in hydrated formula 49.53% 49.53% 49.53%
46.53% 46.53% 46.53% Estimated Total Solids in Hydrated 0.721721
0.721721 0.721721 0.764621 0.764621 0.764621 Formula Delivered to
Cavity (g) Estimated wt. gelatin delivered to 0.12 0.12 0.12 0.16
0.16 0.16 cavity (g) Estimated Water weight delivered 0.71 0.71
0.71 0.67 0.67 0.67 to tooth extraction socket (g) Estimated wt. %
drug-dosed 83.13% 83.13% 83.13% 78.47% 78.47% 78.47% microspheres
in dry formula Estimated wt. % gelatin in dry 16.87% 16.87% 16.87%
21.53% 21.53% 21.53% formula Ratio of water to gelatin 5.82 5.82
5.82 4.04 4.04 4.04 Formulation 7 8 9 10 Target Bupivacaine Dose
over 4-Day 0.300 0.300 0.300 0.300 Period (grams) Estimated Tooth
extraction socket 1.3 1.3 1.3 1.3 Volume (cm.sup.3) Estimated
density of hyrdated 1.1 1.1 1.1 1.1 formula (g/cc) Estimated grams
of hydrated 1.43 1.43 1.43 1.43 formula delivered to tooth
extraction socket (g) Theoretical wt. % Bupivacaine in 50% 50% 50%
50% microspheres Estimated weight of drug-dosed 0.60 0.60 0.60 0.60
microspheres in hydrated formula (g) Estimated % Total Solids in
56.47% 56.47% 56.47% 53.47% Hydrated Formula Wt. % Microspheres
dispersed in 41.96% 41.96% 41.96% 41.96% hydrated formula Wt. %
Gelatin in hydrated formula 14.51% 14.51% 14.51% 11.51% Wt. % Water
in hydrated formula 43.53% 43.53% 43.53% 46.53% Estimated Total
Solids in Hydrated 0.807521 0.807521 0.807521 0.764621 Formula
Delivered to Cavity (g) Estimated wt. gelatin delivered to 0.21
0.21 0.21 0.16 cavity (g) Estimated Water weight delivered 0.62
0.62 0.62 0.67 to tooth extraction socket (g) Estimated wt. %
drug-dosed 74.30% 74.30% 74.30% 78.47% microspheres in dry formula
Estimated wt. % gelatin in dry 25.70% 25.70% 25.70% 21.53% formula
Ratio of water to gelatin 3.00 3.00 3.00 4.04 Formulation 11 12 13
Target Bupivacaine Dose over 4-Day 0.300 0.300 0.000 Period (grams)
Estimated Tooth extraction socket 1.3 1.3 1.3 Volume (cm.sup.3)
Estimated density of hyrdated 1.1 1.1 1.1 formula (g/cc) Estimated
grams of hydrated 1.43 1.43 1.43 formula delivered to tooth
extraction socket (g) Theoretical wt. % Bupivacaine in 50% 50% 50%
microspheres Estimated weight of drug-dosed 0.60 0.60 0.00
microspheres in hydrated formula (g) Estimated % Total Solids in
53.47% 53.47% 25.00% Hydrated Formula Wt. % Microspheres dispersed
in 41.96% 41.96% 0.00% hydrated formula Wt. % Gelatin in hydrated
formula 11.51% 11.51% 25.00% Wt. % Water in hydrated formula 46.53%
46.53% 75.00% Estimated Total Solids in Hydrated 0.764621 0.764621
0.357468345 Formula Delivered to Cavity (g) Estimated wt. gelatin
delivered to 0.16 0.16 0.36 cavity (g) Estimated Water weight
delivered 0.67 0.67 1.07 to tooth extraction socket (g) Estimated
wt. % drug-dosed 78.47% 78.47% 0.00% microspheres in dry formula
Estimated wt. % gelatin in dry 21.53% 21.53% 100.00% formula Ratio
of water to gelatin 4.04 4.04 3.00
[0198] Formula #12 represented an intermediate region in the
statistical design of experiments (DOE) matrix with an intermediate
dry weight % of gelatin binder. A statistically designed experiment
like a Taguchi design can be thought of as a multi-dimensional
exploration space, where the dimensional boundaries of the space
are dictated by the upper and lower factor limits. This space that
is encompassed by the DOE is often referred to as the "design
space." When the statistically designed experiment is executed,
some of the experiments will be executed with a set of factor
values that cause the resulting sample to reside closer to the
middle of the design space than others. Formula #12 is such a
sample. Formula #12 also called for a 4/1 w/w ratio of water to
gelatin binder. After mixing the #12 powdered formula with water,
the hydrated formula #12 was observed to be relatively dry with
significantly lower compliance than a comparable 4/1 w/w water to
neat SF mixture, and lower in compliance than a 3/1 w/w water to
neat SF mixture. Within 12 hours of aging inside of a sealed vial,
formula #12 had become exceedingly stiff and non-compliant. The
formulation was deemed to be too stiff and non-compliant for use in
the end-application. Based on observations taken from the prior
gelatin/water mixing experiments, as described in Parts A and B of
Example 1 above, this time-dependent change in compliance was
likely due to incomplete gelation during the initial mixing process
coupled with a time-dependence that was needed to achieve
equilibrium network formation after mixing.
[0199] In the next step, additional water was added to the hydrated
#12 formula, adjusting the sample to 5.8/1 w/w water to SF in an
attempt to achieve better formability. The initial result was
improved plasticization and higher compliance. This result showed
that with higher levels of water, higher compliance characteristics
were possible, and mechanical efficacy for end use deployment was
indeed possible. However, the addition of more water only served to
dilute the weight percentages of all solids, including active
ingredients like BUP. In a volume restricted application, this
works against the goal of achieving higher bupivacaine dosage
levels.
[0200] In the present example, 1 cc was used as an estimate for the
volume of a tooth extraction socket cavity. Given that the volume
of a tooth extraction socket can vary from
individual-to-individual, even lower socket volumes will be
encountered during end use and deployment. Given that there are
volume-restricted limitations for the hydrated formulation in this
application, it would be even more difficult to achieve higher
bupivacaine dosages with smaller cavity volumes. In order to
illustrate this problem, the estimated cavity volume was reduced to
0.55 cc for a test-set of calculations using the DOE DRAFT-1 factor
constraints.
[0201] Based on a tooth extraction socket volume estimate of about
0.55 cc, Formula #12, comprising about 21% Surgifoam binder on a
dry weight % basis, and the entire DRAFT-1 DOE space was deemed to
be incapable of delivering a targeted bupivacaine dosage of 360 mg.
In fact, based on a 0.55 cc volume, Formula #12 would deliver only
about 150 mg of BUP at best. Substantially higher PLGA microsphere
levels and lower binder levels would be required to increase the
bupivacaine dosage. Given that Surgifoam was found to be a
relatively weak network-forming gel at low concentrations (see Part
A above); and given that even more water would be needed to
plasticize a formula with progressively lower concentrations of
binder, it was hypothesized that SF would not be a suitable binder.
This hypothesis was put to the test as described below.
Part-D. Limitations of High BUP Dosage Delivery Devices with SF as
a Binder
[0202] Using Surgifoam as the binder, the dosage of bupivacaine was
pushed to higher levels. Table 1-6 provides specifications for a
DOE entitled "DRAFT-4," having the same factors as DRAFT-1
described in Part-C above, but with new upper and lower limits.
Table 1-7 provides the wet, hydrated weight percent compositions of
the DOE DRAFT-4 formulas. In conceptualizing DRAFT-4, formula #12
of DRAFT-1 became an upper boundary point in the DOE space with
approximately 21% gelatin binder on a dry weight % basis, referred
to as formula #1 with SF in DOE DRAFT-4. The lowest binder level of
the DRAFT-4 DOE was about 16% gelatin, formula #7 with SF and
formula #7B with GLBG. Table 1-8 provides additional information on
three specific DOE DRAFT-4 formulas that were mixed and
evaluated.
TABLE-US-00006 TABLE 1-6 DOE DRAFT-4 specifications calculations
for creation of dry compositions, and for calculations presented in
Table 1-7. The Factor-3 distribution types were defined as follows,
where PLGA particle sizes were combined according to the following
equation: X = weight fraction of 3.4 micron PLGA particles; Y =
weight fraction of 42.7 micron PLGA particles; 1 - X - Y = weight
fraction of 42.1 micron PLGA particles; type-1 = 0*X with 0*Y;
type-2 = 0.05*X with 0.15*Y; type-3 = 0.1*X with 0.3*Y. FACTOR-3
FACTOR-1 FACTOR-2 PLGA FACTOR-4 Bupivacaine Gelatin micro-sphere
Hydration Expt. dosage (g) Type distribution-type Solution Type 1
0.300 SF 1 pH-neutral 2 0.300 GLBG 2 1% citric 3 0.300 GLPG 3 1%
di-Na citrate 4 0.330 SF 2 1% di-Na citrate 5 0.330 GLBG 3
pH-neutral 6 0.330 GLPG 1 1% citric 7 0.360 SF 3 1% citric 8 0.360
GLBG 1 1% di-Na citrate 9 0.360 GLPG 2 pH-neutral 10 0.300 SF 100%
3.4 um pH-neutral 11 0.300 SF 100% 42.7 um pH-neutral 12 0.300 SF 1
pH-neutral 13 0.000 SF No PLGA pH-neutral .sup. 7B 0.360 GLBG 3 1%
citric
TABLE-US-00007 TABLE 1-7 Hydrated) weight % compositions for DOE
DRAFT-4. Wt. % PLGA Expt. Wt. % gelatin Microspheres Wt. % water
solution 1 11.51% 41.96% 46.53% 2 11.51% 41.96% 46.53% 3 11.51%
41.96% 46.53% 4 10.68% 46.15% 43.16% 5 10.68% 46.15% 43.16% 6
10.68% 46.15% 43.16% 7 8.70% 44.46% 46.84% 8 9.85% 50.35% 39.80% 9
9.85% 50.35% 39.80% 10 11.51% 41.96% 46.53% 11 11.51% 41.96% 46.53%
12 9.05% 32.99% 57.95% 13 19.84% 0.00% 80.16% .sup. 7B 8.70% 44.46%
46.84%
TABLE-US-00008 TABLE 1-8 DOE DRAFT-4 calculations of pertinent
composition information based on the constraints presented in Table
1-6. Although placebo PLGA microspheres were used in preparing
samples, calculations were performed to estimate a theoretical
dosage of BUP delivery to a tooth extraction socket of 1.3 cc
volume, assuming that the PLGA microspheres were loaded with 50%
BUP by weight. Formulation 1 7 7B Target Bupivacaine Dose over
4-Day Period (grams) 0.300 0.360 0.360 Estimated Tooth extraction
socket Volume (cm.sup.3) 1.3 1.3 1.3 Estimated density of hydrated
formula (g/cc) 1.1 1.1 1.1 Estimated grams of hydrated formula
delivered to tooth 1.43 1.43 1.43 extraction socket (g) Theoretical
wt. % Bupivacaine in microspheres 50% 50% 50% Estimated weight of
drug-dosed microspheres in hydrated 0.60 0.72 0.72 formula (g)
Estimated % Total Solids in Hydrated Formula 53.47% 60.20% 60.20%
Wt. % Microspheres dispersed in hydrated formula 41.96% 50.35%
50.35% Wt. % Gelatin in hydrated formula 11.51% 9.85% 9.85% Wt. %
Water in hydrated formula 46.53% 39.80% 39.80% Estimated Total
Solids in Hydrated Formula Delivered to Cavity 0.764621 0.86087619
0.860876 (g) Estimated wt. gelatin delivered to cavity (g) 0.16
0.14 0.14 Estimated Water weight delivered to tooth extraction
socket (g) 0.67 0.57 0.57 Estimated wt. % drug-dosed microspheres
in dry formula 78.47% 83.64% 83.64% Estimated wt. % gelatin in dry
formula 21.53% 16.36% 16.36% Ratio of water to gelatin 4.04 4.04
4.04 Binder Type SF SF GLBG Water Type pH-neutral 1% citric 1%
citric Sphere distribution type 1 3 3
[0203] Upon mixing formula #7 with Surgifoam as the binder, the
resulting device was observed to be a non-compliant dry-blend at a
4 to 1 water/binder w/w ratio. Addition of more water for
plasticization was fruitless owing to weakening of the binder
network. Upon mixing formula #1 with an even higher SF binder
level, the resulting device was also observed to be a non-compliant
dry-blend at a 4 to 1 water/binder w/w ratio. Thus, with Surgifoam
gelatin appearing to be an unacceptable binder at low binder levels
and at higher water/binder w/w ratios, a new analogous formula was
made with the Great Lakes bovine gelatin (formula 7B). Although
neat GLBG was previously observed to be slower to gel and slower to
reach equilibrium compliance than neat SF (see Part-A above), it
was also noted to become a stronger and more elastic gel than SF
when plasticized with water at equivalent water to gelatin weight
ratios. The ability of GLBG to form a stronger gelled network than
SF was hypothesized to be a possible solution to the problem of
trying to balance the need for achieving acceptable composite
properties at the low binder levels and at the higher volume
fractions of microspheres that are needed for delivery of higher
bupivacaine dosages to small fixed-volume cavities. Indeed, upon
mixing with water, Formula #7B was observed to congeal to form a
dough-like material at a 4 to 1 water/binder weight ratio. The
compliance of Formula #7B was still relatively low, but unlike
Surgifoam, Formula #7B had nevertheless congealed to form a
compliant solid, which was not a flaky dry-blend.
[0204] Thus, at the low binder levels necessitated by volume
restrictions and by elevated BUP target delivery dosages, the
preferred binder is one that is strong enough to provide acceptable
cohesive integrity, and it is also one that has the ability to
provide acceptable gel-network formation at relatively low
concentrations. In this regard, even though neat GLBG is less
compliant than neat SF when plasticized with equivalent levels of
water, a more elastic, lower-compliance gelatin such as GLBG is
preferred as a binder over Surgifoam when it is used in a composite
mixture containing PLGA microspheres dispersed in a hydrated
gelatin matrix.
[0205] Hydrated formula #7B became increasingly stiff and lower in
compliance as it was aged in a closed container, consistent with
the time-dependent changes in rheological characteristics that were
observed in the prior experiments with gelatin and water alone.
Thus, GLBG, like Surgifoam, was observed to still require higher
levels of water for plasticization. Again, although GLBG was deemed
to be a better binder than SF, this is not a desirable direction
for achieving higher bupivacaine dosages in a fixed volume end use
application. These results also showed that network formation would
be time-dependent, and that equilibrium conditions might require
several hours or more at any given water-level.
Example 2. Design of a Controlled Release Device for Delivering BUP
within a Volume-Restricted End Use Application
[0206] A statistically designed experiment entitled "DOE DRAFT-6"
was constructed to demonstrate the limitations of an embodiment
whereby dry powders of bupivacaine-loaded PLGA microspheres and
gelatin would be pre-masticated with water and then delivered as a
compliant dough-like material during end use. Given the limitations
of SF as demonstrated in Example 1, the gelatins of choice for this
example were GLBG and GLPG.] Note that neat SF gels are routinely
used by clinicians as hemostats to fill the tooth extraction socket
in post tooth extraction applications. As such, neat SF gels are
recognized by clinicians as having reasonably acceptable mechanical
compliance and formability characteristics. For this reason, neat
SF gels were used as qualitative benchmarks for targeting
acceptable compliance characteristics, while simultaneously
attempting to maximize the theoretical BUP dosage delivery
limits.
[0207] The goal of the DOE DRAFT-6 experiment was to create a
viable mixture that could be used for the following purposes: 1)
for use in analytical testing to investigate and develop desirable
bupivacaine time-release profiles; 2) to simultaneously provide
mechanical compliance characteristics similar to what many
clinicians would recognize as an acceptable benchmark similar to
neat SF gelled with water; 3) to simultaneously maximize the
theoretical dosage limit of BUP while working with restrictions
presented by a fixed volume constraint; and 4) to produce a viable
end use formulation that can be pre-hydrated with water before
deployment to deliver relatively high dosages of BUP in a
volume-restricted end use application.
[0208] Considerations for the conceptualization and creation of DOE
DRAFT-6 can be summarized as follows: [0209] (1) the volume
restriction for the end-application, estimated to be ca. 0.55 cc in
this example, causes the upper limit dosage of bupivacaine to be
severely constrained. In order to retain mechanical efficacy,
compliance and formability, there is a need for some minimum level
of binder and water, which places a limitation on the maximum
weight % concentration of PLGA microspheres that can be
incorporated into the device for use and deployment in a
volume-restricted environment; [0210] (2) higher levels of
bupivacaine loading in the PLGA microspheres would be required to
reach bupivacaine delivery dosages of >60 mg. A level of 20% w/w
BUP in the PLGA microspheres would lead to maximum dose deliveries
of less than 60 mg; [0211] (3) lower binder levels would be
required to maximize the microsphere content and hence to maximize
the bupivacaine delivery dosage. This is a constraint that weakens
the composite and necessitates not only the use of better
network-forming binders, but also higher levels of volume-occupying
water for plasticization; [0212] (4) lower binder levels
necessitate higher molecular weight network-forming gels that are
susceptible to time-dependent reductions in compliance owing to
diffusion-rate limitations which impact the time required for the
gelling network to reach its equilibrium state; and [0213] (5)
diffusion rates and time-dependent compliance characteristics are
further confounded by both the particle size distribution of the
microspheres, which also affects the bupivacaine time-release
profile, and by the particle sizes of the gelatin particulates.
Other considerations included identifying the controlled factors
for DOE DRAFT-6 and determining the boundary limits for factors. A
general description of the factors and considerations pertaining to
boundary limits are described here: [0214] Factor-1: the weight %
binder range was chosen to help produce a potentially viable device
while maximizing the bupivacaine dosage within the limitations of
the embodiment; [0215] Factor-2: the gelatin type, where GLBG and
GLPG and mixtures thereof were chosen to help achieve acceptable
mechanical properties at the low binder levels necessitated by the
desire to achieve higher bupivacaine dosages with the 0.55 cc
volume constraint; [0216] Factor-3: the microsphere distribution
type, where different particle size distributions would lead to
different surface-to-volume ratios for the purposes of impacting
mechanical properties, and for the purpose of modulating
bupivacaine release and diffusion rates; and [0217] Factor-4: the
use of pH modulators, which affect gel-rate and gel strength. The
pH modulators are also anticipated to affect bupivacaine free-base
solubility, bupivacaine release rate, lactic acid formation rate,
and lactate neutralization.
[0218] The DOE factors and levels for DRAFT-6 are provided in Table
2-1. A Taguchi 4-factor, 3-level design template was employed,
represented by experiments 1 through 9 in Table 2-1, along with
four additional one-off experiments, 10 through 13, where #13
represented a 4.04/1 w/w water to SF benchmark.
[0219] The statistical DOE factors included: 1) the weight % of
gelatin binder on a dry-basis, wherein the range was chosen so as
to produce a potentially viable device while maximizing BUP dosage
(21.53%, 18.80%, 16.36%); 2) the gelatin type (GLBG, GLPG, and a
50/50 w/w mixture of the two); 3) microsphere particle size
distribution-type using mixtures of the PLGA microspheres from SWRI
that were described in Example 1 (distribution type-1i=100% D50
42.1 micron emulsion particles; distribution type-2=80% D50 42.1
micron emulsion particles+15% D50 42.7 micron spinning-disc
particles+5% D50 3.4 micron spinning-disc particles; distribution
type-3=60% D50 42.1 micron emulsion particles+30% D50 42.7 micron
spinning-disc particles+10% D50 3.4 micron spinning-disc
particles); and 4) pH modulators incorporated into solutions with
distilled water for hydrating the dry powder mixtures (pH-neutral
water, 1% citric acid solution, and 1% di-sodium citrate solution)
which have been shown to affect gel-rate and gel strength and are
anticipated to affect BUP free-base solubility, BUP release rate,
lactic acid formation rate, and lactate neutralization.
[0220] For the purposes of this example, the hydrated formula
mixture weights were targeted to be between 0.7 g and 0.9 g. The
initial water to gelatin binder ratio was specified to be 4.04/1
w/w. Calculations depicting various attributes of DOE DRAFT-6 are
provided in Tables 2-2, 2-3, and 2-4, respectively. With the
restraint that the cavity volume was estimated to be 0.55 cc in
this example, the bupivacaine delivery dosages were limited to
those as described in Table 2-5.
[0221] Samples 10, 11, and 12 comprising segregated particle
distributions were mixed first, followed by the statistical
DOE-space samples 1 through 9. Whenever possible, qualitative
trends and observations were noted immediately after mixing. Given
that it takes time for the networks to reach equilibrium, the
hydrated samples made with a weight ratio of water to gelatin of
4.04/1 w/w were placed into closed containers for 24 hours. The
hydrated samples were then removed and were qualitatively ranked
for their relative compliance, for their relative degree of
re-formability, and for their relative tackiness during handling.
In the next step, a small amount of additional water was added to
rehydrate the samples. The added level of water resulted in an
increase in the total weight ratio of water to gelatin binder from
an initial value of 4.04/1 w/w to a value of 5.54/1 w/w. This had
the effect of diluting the fixed-volume compositions and allowed
for qualitative ranking of relative cohesive strength after aging.
The rehydrated compositions are provided in Table 2-5.
[0222] Whenever possible, the qualitative rankings were used as
responses for trend analyses, and for determining the significance
of the controlled factors. Design-Ease 9 DOE software (Stat-Ease,
Inc.) was used to test for significance of differences at the 95%
confidence level (CL).
TABLE-US-00009 TABLE 2-1 DOE DRAFT-6 specifications calculations
for creation of dry compositions, and for calculations presented in
Tables 2-2, 2-3, 2-4, and 2-5. The Factor-3 distribution types were
defined as follows, where PLGA particle sizes were combined
according to the following equation: X = weight fraction of 3.4
micron PLGA particles; Y = weight fraction of 42.7 micron PLGA
particles; 1 - X - Y = weight fraction of 42.1 micron PLGA
particles; type-1 = 0*X with 0*Y; type-2 = 0.05*X with 0.15*Y;
type-3 = 0.1*X with 0.3*Y. FACTOR-1 wt. % FACTOR-3 FACTOR-4 gelatin
PLGA micro- Hydration in dry FACTOR-2 sphere distri- Solution Expt.
formula Gelatin Type bution-type Type 1 21.53% GLBG 1 pH-neutral 2
21.53% GLPG 2 1% citric 3 21.53% 50/50 3 1% di-Na GLBG/GLPG citrate
4 18.80% GLBG 2 1% di-Na citrate 5 18.80% GLPG 3 pH-neutral 6
18.80% 50/50 1 1% citric GLBG/GLPG 7 16.36% GLBG 3 1% citric 8
16.36% GLPG 1 1% di-Na citrate 9 16.36% 50/50 2 pH-neutral
GLBG/GLPG 10 18.80% GLBG 100% 3.4 .mu.m pH-neutral 11 18.80% GLBG
100% 42.7 .mu.m pH-neutral 12 18.80% GLBG 1 pH-neutral 13 100.00%
SF No PLGA pH-neutral
TABLE-US-00010 TABLE 2-2 Dry weight % compositions for DOE DRAFT-6.
Wt. % PLGA Expt. Wt. % gelatin Microspheres 1 21.5% 78.5% 2 21.5%
78.5% 3 21.5% 78.5% 4 18.8% 81.2% 5 18.8% 81.2% 6 18.8% 81.2% 7
16.4% 83.6% 8 16.4% 83.6% 9 16.4% 83.6% 10 18.8% 81.2% 11 18.8%
81.2% 12 18.8% 81.2% 13 100.0% 0.0%
TABLE-US-00011 TABLE 2-3 Wet (hydrated) weight % compositions for
DOE DRAFT-6; water to gelatin w/w ratio was specified to be 4.04/1.
Wt. % PLGA Expt. Wt. % gelatin Microspheres Wt. % water solution 1
11.51% 41.96% 46.53% 2 11.51% 41.96% 46.53% 3 11.51% 41.96% 46.53%
4 10.69% 46.15% 43.16% 5 10.69% 46.15% 43.16% 6 10.69% 46.15%
43.16% 7 9.85% 50.35% 39.80% 8 9.85% 50.35% 39.80% 9 9.85% 50.35%
39.80% 10 10.69% 46.15% 43.16% 11 10.69% 46.15% 43.16% 12 10.69%
46.15% 43.16% 13 19.84% 0.00% 80.16%
TABLE-US-00012 TABLE 2-4 Wet (re-hydrated) weight % compositions
for DOE DRAFT-6; water to gelatin w/w ratio was 5.54/1. Wt. % PLGA
Expt. Wt. % gelatin Microspheres Wt. % water solution 1 9.82%
35.78% 54.40% 2 9.82% 35.78% 54.40% 3 9.82% 35.78% 54.40% 4 9.21%
39.78% 51.01% 5 9.21% 39.78% 51.01% 6 9.21% 39.79% 51.00% 7 8.58%
43.87% 47.55% 8 8.58% 43.87% 47.55% 9 8.58% 43.87% 47.55% 10 9.21%
39.78% 51.01% 11 9.21% 39.78% 51.01% 12 9.21% 39.78% 51.01%
TABLE-US-00013 TABLE 2-5 DOE DRAFT-6 calculations of pertinent
composition information based on the constraints presented in Table
2-1. Although placebo PLGA microspheres were used in preparing
samples, calculations were performed to estimate a theoretical
dosage of BUP delivery to a tooth extraction socket of 0.55 cc
volume, assuming that the PLGA microspheres were loaded with 50%
BUP by weight. Formulation 1 2 3 Target Bupivacaine Dose over 4-Day
0.127 0.127 0.127 Period (grams) Estimated Tooth extraction socket
0.55 0.55 0.55 Volume (cm.sup.3) Estimated density of hydrated 1.1
1.1 1.1 formula (g/cc) Estimated grams of hydrated 0.605 0.605
0.605 formula delivered to tooth extraction socket (g) Theoretical
wt. % Bupivacaine in 50% 50% 50% microspheres Estimated weight of
drug-dosed 0.25 0.25 0.25 microspheres in hydrated formula (g)
Estimated % Total Solids in 53.47% 53.47% 53.47% Hydrated Formula
Wt. % Microspheres dispersed in 41.96% 41.96% 41.96% hydrated
formula Wt. % Gelatin in hydrated formula 11.51% 11.51% 11.51% Wt.
% Water in hydrated formula 46.53% 46.53% 46.53% Estimated Total
Solids in Hydrated 0.323499306 0.323499306 0.323499306 Formula
Delivered to Cavity (g) Estimated wt. gelatin delivered to 0.07
0.07 0.07 cavity (g) Estimated Water weight delivered 0.28 0.28
0.28 to tooth extraction socket (g) Estimated wt. % drug-dosed
78.47% 78.47% 78.47% microspheres in dry formula Estimated wt. %
gelatin in dry 21.53% 21.53% 21.53% formula Ratio of water to
gelatin 4.04 4.04 4.04 Binder Type GLBG GLPG 50/50 GLBG/GLPG Water
Type pH-neutral 1% 1% citric diNaCitrate Sphere distribution type 1
2 3 Formulation 4 5 6 Target Bupivacaine Dose over 4-Day 0.140
0.140 0.140 Period (grams) Estimated Tooth extraction socket 0.55
0.55 0.55 Volume (cm.sup.3) Estimated density of hydrated 1.1 1.1
1.1 formula (g/cc) Estimated grams of hydrated 0.605 0.605 0.605
formula delivered to tooth extraction socket (g) Theoretical wt. %
Bupivacaine in 50% 50% 50% microspheres Estimated weight of
drug-dosed 0.28 0.28 0.28 microspheres in hydrated formula (g)
Estimated % Total Solids in 56.84% 56.84% 56.84% Hydrated Formula
Wt. % Microspheres dispersed in 46.15% 46.15% 46.15% hydrated
formula Wt. % Gelatin in hydrated formula 10.69% 10.69% 10.69% Wt.
% Water in hydrated formula 43.16% 43.16% 43.16% Estimated Total
Solids in Hydrated 0.34386857 0.34386857 0.34386857 Formula
Delivered to Cavity (g) Estimated wt. gelatin delivered to 0.06
0.06 0.06 cavity (g) Estimated Water weight delivered 0.26 0.26
0.26 to tooth extraction socket (g) Estimated wt. % drug-dosed
81.20% 81.20% 81.20% microspheres in dry formula Estimated wt. %
gelatin in dry 18.80% 18.80% 18.80% formula Ratio of water to
gelatin 4.04 4.04 4.04 Binder Type GLBG GLPG 50/50 GLBG/GLPG Water
Type 1% pH-neutral 1% diNaCitrate citric Sphere distribution type 2
3 1 Formulation 7 8 9 Target Bupivacaine Dose over 4-Day 0.152
0.152 0.152 Period (grams) Estimated Tooth extraction socket 0.55
0.55 0.55 Volume (cm.sup.3) Estimated density of hydrated 1.1 1.1
1.1 formula (g/cc) Estimated grams of hydrated 0.605 0.605 0.605
formula delivered to tooth extraction socket (g) Theoretical wt. %
Bupivacaine in 50% 50% 50% microspheres Estimated weight of
drug-dosed 0.30 0.30 0.30 microspheres in hydrated formula (g)
Estimated % Total Solids in 60.20% 60.20% 60.20% Hydrated Formula
Wt. % Microspheres dispersed in 50.35% 50.35% 50.35% hydrated
formula Wt. % Gelatin in hydrated formula 9.85% 9.85% 9.85% Wt. %
Water in hydrated formula 39.80% 39.80% 39.80% Estimated Total
Solids in Hydrated 0.364216852 0.364216852 0.364216852 Formula
Delivered to Cavity (g) Estimated wt. gelatin delivered to 0.06
0.06 0.06 cavity (g) Estimated Water weight delivered 0.24 0.24
0.24 to tooth extraction socket (g) Estimated wt. % drug-dosed
83.64% 83.64% 83.64% microspheres in dry formula Estimated wt. %
gelatin in dry 16.36% 16.36% 16.36% formula Ratio of water to
gelatin 4.04 4.04 4.04 Binder Type GLBG GLPG 50/50 GLBG/GLPG Water
Type 1% 1% pH-neutral citric diNaCitrate Sphere distribution type 3
1 2 Formulation 10 11 12 Target Bupivacaine Dose over 4-Day 0.140
0.140 0.140 Period (grams) Estimated Tooth extraction socket 0.55
0.55 0.55 Volume (cm.sup.3) Estimated density of hydrated 1.1 1.1
1.1 formula (g/cc) Estimated grams of hydrated 0.605 0.605 0.605
formula delivered to tooth extraction socket (g) Theoretical wt. %
Bupivacaine in 50% 50% 50% microspheres Estimated weight of
drug-dosed 0.28 0.28 0.28 microspheres in hydrated formula (g)
Estimated % Total Solids in 56.84% 56.84% 56.84% Hydrated Formula
Wt. % Microspheres dispersed in 46.15% 46.15% 46.15% hydrated
formula Wt. % Gelatin in hydrated formula 10.69% 10.69% 10.69% Wt.
% Water in hydrated formula 43.16% 43.16% 43.16% Estimated Total
Solids in Hydrated 0.34386857 0.34386857 0.34386857 Formula
Delivered to Cavity (g) Estimated wt. gelatin delivered to 0.06
0.06 0.06 cavity (g) Estimated Water weight delivered 0.26 0.26
0.26 to tooth extraction socket (g) Estimated wt. % drug-dosed
81.20% 81.20% 81.20% microspheres in dry formula Estimated wt. %
gelatin in dry 18.80% 18.80% 18.80% formula Ratio of water to
gelatin 4.04 4.04 4.04 Binder Type GLBG GLBG GLBG Water Type
pH-neutral pH-neutral pH-neutral Sphere distribution type 3.4
micron 42.7 micron 42.1 micron
[0223] Qualitative trend analysis after initial mixing of samples
10, 11, and 12 revealed that the highest cohesive strength was
achieved in the sample made with the D50=3.4 micron microspheres,
followed by the sample made with the D50=42.1 micron microspheres.
This trend also seemed to manifest itself among the DOE-space
samples 1-9. Samples with the highest fraction of 3.4 micron
particles trended towards displaying the best cohesive strength
after mixing. This result indicates that from a mechanical property
perspective, it is desirable to maximize the smaller particle size
particle fraction while simultaneously balancing the overall
distribution to achieve the desired bupivacaine release profile,
particularly since smaller particles will release BUP faster than
larger ones owing to their higher surface to volume ratio.
[0224] The 42.1 microsphere particles led to expedient mixing with
minimal clumping when compared to their 42.7 micron
spinning-disc-dried counterparts. It is noted that the 42.7 micron
spinning-disc-dried particles were also agglomerated, whereas the
42.1 micron emulsion particles were more free-flowing.
[0225] Qualitative compliance was observed to increase with
increasing gelatin binder level, as expected, and with fewer
microspheres.
[0226] Because the samples were initially evaluated while they were
in a dynamic state before they had reached their time-dependent
equilibrium properties, the samples were placed in closed
containers and were then allowed to equilibrate for 24 hours.
Qualitative trend analyses at t=24 hours after mixing showed that
each of the samples had increased in stiffness, a result which was
consistent with the earlier DRAFT-1 DOE results of Example 1.
Sample #10, which was made exclusively with 3.4 micron particles,
continued to exhibit higher cohesive strength characteristics than
any of the other samples.
[0227] Statistical trend analyses of the categoric factors from the
DOE space indicated that the qualitative relative compliance
response at t=24 hours after mixing with water was significantly
affected by the weight % binder and by the binder-type at the 95%
confidence level (CL), with p values <0.05, and with higher
binder leading to higher compliance. These results showed that the
minimum tolerable threshold for the binder level is between 21.5%
and 18.8% by weight of the dry formula, with GLBG bovine gelatin
being the preferred binder. Of course, higher levels of binder and
water would always be helpful from a mechanical property
perspective, but this would be counter to the objective of
developing a formula with maximal BUP dosage potential.
[0228] Statistical trend analyses also revealed that the relative
tack response characteristics scaled significantly with the weight
% binder at the 95% CL. Within the DOE space, the minimum binder
threshold for achieving the best tack appeared to be at or near
about 19% by weight of the dry formula (p<0.05). This result
reaffirms that for the purpose of creating a powder-based formula,
the dry binder level should be maximized to a level of greater than
about 18%. Even higher levels would be desirable, but only to the
degree that lower bupivacaine delivery dosages can be tolerated in
the application.
[0229] After rehydration with additional water, each of the samples
was observed to exhibit an increase in relative compliance. Again,
sample #10, which was made exclusively with 3.4-micron particles,
was unique in that it exhibited the best physical properties.
Specifically, sample #10 exhibited the highest relative cohesive
strength and homogeneity of all the samples. This observation was
consistent with the DOE-space trend analyses. Namely, the relative
cohesive strength characteristics for the rehydrated formulas were
qualitatively observed to increase as the percentage of 3.4 micron
particles was increased within the formulation by employing sphere
distribution type-2 or type-3 as depicted in Table 2-1, wherein
type-2 equates to the particle distribution being comprised of 5%
D50=3.4 micron particles and type-3 equates to the particle size
distribution being comprised of 10% D50=3.4 micron particles vs.
type-1 which contains no added D50=3.4 micron particles.
[0230] The results suggest that the dispersed PLGA particles
augment the mechanical properties of the hydrated gelatin network.
In other words, the PLGA particles do not simply behave as
dispersed filler particles which deteriorate mechanical properties
or provide no improvement. Instead, they behave as reinforcing
fillers which improve mechanical properties. This means that they
not only perform a primary function of encapsulating active
ingredients for controlled-release, they also perform a beneficial
secondary function of reinforcing the hydrated binder matrix, with
smaller PLGA particle sizes having a more pronounced positive
effect. This further implies that the PLGA microspheres will not
only provide a first diffusion barrier for the release of BUP or
other active ingredients, but its reinforcing presence in the
matrix will also affect the compliance of the hydrated gelatin
polymer itself, which in turn will further augment diffusion rates
of BUP through the gelled matrix phase once the BUP has already
diffused from the dispersed PLGA microspheres and into the gelled
matrix. Also, from a macroscopic perspective, the mechanical
reinforcement of the gelled gelatin binder by PLGA particles will
also increase the resistance to erosion of the formulation within
the end use application.
[0231] Trend analyses of the rehydrated samples also revealed a
moderately detectable effect of pH-modulator on cohesive strength
after re-hydration (p-value.about.0.10), with citric acid having a
positive effect on cohesive strength and with di-sodium citrate
having no detectable effect. Although these trends were not as
significant at the 95% CL as other trends, they were nevertheless
reasonable, particularly in light of the other qualitative findings
that were presented in Example 1. Namely, the presence of citric
acid was shown in Example 1 to lead to an increase in gelation
rates for the neat proteins. Indeed, this trend seemed to manifest
itself even when the gelatin proteins were used as binders in
samples containing dispersed PLGA microspheres.
[0232] Based on the collective set of DOE responses, an embodiment
of a delivery system using a formulation comprising a powdered
mixture appears restricted to deliver a dosage of no more than
about 140 mg bupivacaine to a 0.55 cc cavity, and only then by
assuming that the % bupivacaine loading in the PLGA microspheres is
increased from 20% to 50% by weight. Low gelatin binder levels are
also required to maximize the volume fraction of microspheres and
bupivacaine. It appears that the lower limit threshold for the
binder is approximately 18% of the dry weight. At these relatively
low levels, a network-forming gelatin like GLBG (Bloom=225 g) is
preferred for its ability to impart the type of cohesive strength
that is needed to bind the spheres together when the device is
hydrated.
[0233] If the product is intended to be premixed with water, and if
higher bupivacaine dosages are desired, then the occupied volume of
water must also be accounted for, and the water-level should be
minimized since it will effectively dilute the microsphere
concentration and will further reduce the maximum bupivacaine
delivery dosage to levels less than 140 mg if too much water is
employed.
[0234] For reasons pertaining to mechanical properties, it is also
preferable to skew the PLGA particle size distribution towards
smaller particles, but only to the degree that this can be
tolerated depending on bupivacaine time-release profile
targets.
[0235] Larger PLGA microspheres made via an emulsion process
provide qualitatively lower formula viscosities than their
spinning-disc/spray-dried counterparts. In essence, this equates to
a higher PLGA loading potential during mixing, which is also
directionally preferred for achieving higher bupivacaine dosages,
but only to the degree that adequate compliance and cohesive
strength can be maintained. The D50=42.1 micron emulsion particles
were also observed to mix more uniformly with faster wetting than
their similarly-sized spinning-disc spray-dried counterparts, the
D50=42.7 micron placebo PLGA microspheres. The emulsion particles
(42.1 um) are thus preferred for the present application to the
degree that larger particles are needed to achieve targeted release
profiles.
[0236] Again, smaller particles (D50=3.4 microns) made by
spinning-disc methods, by spray-drying with spinning disc, or by
emulsion processes are desirable for reasons pertaining to
mechanical properties, but only to the degree that their higher
surface-to-volume ratios and release characteristics can be
conducive to achieving specific time-dependent bupivacaine release
profile targets.
[0237] Although release profile targets will be end use specific,
it should be appreciated from these teachings that there will be
several adjustable factors besides PLGA surface-to-volume ratios
that can also conceivably be used to modulate and control the
time-release profiles of bupivacaine and the like. For example,
citric acid (a Bronsted acid) or di-sodium citrate (a Bronsted
base) was observed to be viable with no obvious deleterious effects
on rheology or properties of the delivery system. Citric acid was
observed to enhance binder network formation. From this
perspective, these types of compounds can serve dual functions.
They can be used to modulate the physical properties of the binder
system, and their activity can also be exploited for the dual
purpose of modulating the solubility of the bupivacaine free
base.
[0238] For example, a Bronsted acid will enhance the solubility of
bupivacaine free base as it is released from a PLGA particle,
thereby enhancing its bioavailability. Conversely, a Bronsted base
would skew the acid-base equilibrium towards more bupivacaine free
base, thereby reducing its bioavailability. Further, these types of
compounds can be employed directly as powdered ingredients, which
would make them immediately available upon hydration of the device.
In addition, these types of compounds can be optionally and
separately microencapsulated, which would attenuate their
availability for acid-base interactions with bupivacaine, either
with bupivacaine's acidic form or its free-base form.
[0239] By balancing these various types of formulation levers in
combination, use of citric acid and use of a gelatin with a higher
Bloom value like GLBG, it can be appreciated that one could achieve
targeted bupivacaine release profiles while simultaneously
employing higher fractions of high surface-to-volume particles if
so desired. For example, with the combined use of these levers, one
could potentially use a higher fraction of 3.4 micron PLGA
particles than would otherwise be viable. Again, this direction
might be desirable for reasons pertaining to achieving improved
mechanical properties, which in turn could be leveraged to achieve
lower net binder levels and higher net PLGA levels with higher net
bupivacaine dosages.
[0240] Based on the above results, a mixed-particle size
distribution delivery system to evaluate bupivacaine release
profiles would use the particle size distribution of Formula #7,
and would employ the GLBG binder at the levels used in Formulas #4,
#5, and #6. The water-level required for pre-hydration should be
minimized since adding more water equates to bupivacaine dilution.
Furthermore, if the bupivacaine's release character can be
adequately controlled, it would also be desirable to employ citric
acid in the water phase at a concentration of 1% by weight or
higher. The fraction of small particles should then be increased to
the degree permitted based on the targeted bupivacaine release
profiles. It is also preferable to increase the binder level to the
degree permitted based on the target bupivacaine dosage and based
on the required level of liquid water volume that is needed to
achieve the desired compliance for any particular end use.
Example 3. Testing Surgifoam Gelatin as a Binder Component for Use
in a Formulation with Mineral Oil
[0241] In a first step, 0.1 g of Surgifoam was weighed into a small
beaker. In order to batch a formula analogous to #7, #8, or #9 from
DOE DRAFT-6 in Example 2, one would have to add 0.4039 g of water
to the SF to achieve a liquid to gelatin weight ratio=4.04/1 w/w.
However, since the goal of this example is to reduce the volume
fraction of non-PLGA components to facilitate higher PLGA
microsphere and BUP concentrations in the final formulation, the
dry SF would have to be mixed with less than this amount of liquid.
From a compliance perspective, this would be directionally
incorrect if water were to be chosen as the liquid. However, if a
different type of liquid were to be chosen, such as one that had
the ability to simply disperse the gelatin particles without
diffusing into the particles and without prematurely gelling the
particles, then the concept of using less liquid might become more
plausible. In this example, mineral oil (MO) was chosen as the
liquid in place of water (Aldrich Heavy wt. CAS 8020-83-5, product
33,076-0).
[0242] In step 2, 0.1055 g mineral oil was added to 0.1 g SF, but
it formed a dry blend.
[0243] In a third step, more mineral oil was added to bring the net
addition to 0.1524 g. The SF powder began to consolidate into an
array of surface-wetted particles, but the blend was still too dry
and had very little cohesive integrity and could not be pressed or
formed into a shape.
[0244] In step 4, more mineral oil was added, bringing the total to
0.2074 g. Again, it was noted that more oil would still be needed
to form a compliant dispersion/mixture.
[0245] In step 5, the total oil level was increased to 0.3044 g.
The blend was continuing to consolidate and pack into a weak
amalgam, but it was still too dry and too cohesively weak to form a
compliant mixture/dispersion. Based on this result, the approach of
using MO with Surgifoam was abandoned because the objective was to
minimize non-PLGA components while still maintaining sufficient
compliance and cohesive strength to facilitate fibrous
textile-impregnation. It was reasoned that a gelatin binder with a
larger average particle size might produce a liquid dispersion with
less oil, while still providing enough cohesive strength and
compliance for subsequent textile impregnation.
Example 4. Testing GLBG as a Binder Component for Use in a
Formulation with Mineral Oil (Formula #14A)
[0246] In a first step, 0.1054 g MO was added to 0.1022 g GLBG.
Owing to the larger particle size of the GLBG, a completely wet and
flowable/compliant amalgam was formed with only a 1/1 w/w ratio of
liquid oil to gelatin. Thus, in order to maximize the % solids
while simultaneously minimizing the % liquid in the device formula,
and in order to simultaneously provide a hydrophilic binder
component (e.g., GLBG) capable of binding PLGA spheres upon
hydration, this result shows that it is desirable to increase the
particle size of the ground gelatin component, and even to maximize
the gelatin particle size to the degree permissible by the end use
application.
[0247] In the next step, 0.1328 g of the 3.4-micron PLGA
microspheres, and 0.2991 g of the 42.7-micron PLGA microspheres
were weighed into a separate small beaker, approximately 70/30 w/w
large to small PLGA particles, consistent with Example 2. At this
point, the mix was consolidated with a small spatula into a dry
cake. This mixture contained approximately 83.5 weight % total
solids dispersed in mineral oil, capable of delivering
approximately 206 mg bupivacaine to a 0.55 cc cavity. This is
listed as Formula #14A in Table 5-1. It is noted that if Formula
#14A were able to hydrate in vivo, then 14A could also be useful
without fiber reinforcement.
Example 5. Preparation of a Controlled-Release Delivery Formulation
Comprising a Mixture for Stand-Alone Use or for Optional
Impregnation into a Cellulose Fiber Textile (Formula #14B)
[0248] A sample mixture, Formula #14B, analogous to Formula #14A
was prepared with the use of additional mineral oil (MO) for the
purpose of insuring that the mixture could be easily pressed and
impregnated into a cellulose fiber textile to form a reinforced
composite-like structure. The ratio of large to small PLGA
particles in this example was chosen to be 70/30 (w/w). This ratio
was chosen based on results presented in Example 2 above, wherein
the use of higher fractions of small PLGA particles was determined
to be preferred for achieving suitable cohesive strength
characteristics for hydrated devices. Spin-disc spray-dried
42.7-micron microparticles were used in this example to demonstrate
the concept.
[0249] Initially, 0.0985 g of additional mineral oil (MO) was added
to Formula #14A from Example 4, bringing the total level to 0.2039
g mineral oil. At this point the amalgam became a tacky paste. In
spite of having a lower weight percentage of liquid carrier,
Formula #14B had a lower relative viscosity than analogous formulas
from Example 2 that were made with water as the liquid carrier.
Specifically, formulas #1, #2, and #3 in DOE DRAFT-6 each contained
approximately 53% solids by weight with water as the liquid
carrier. By contrast, formula #14B was comprised of 72.36% solids
by weight with oil as the carrier. Thus, by substituting oil for
water as the liquid carrier, it was discovered that a compliant
vehicle could be formed with a higher weight percentage solids.
Consequently, Formula #14B was estimated to be capable of
delivering 177 mg bupivacaine to a 0.55 cc tooth extraction socket
assuming a 50% w/w loading of BUP in the PLGA microspheres as shown
in Table 5-1. By contrast, as previously noted in Table 2-5,
comparable formulas with water as the carrier were only capable of
delivering BUP dosages of 127 to 150 mg at best.
[0250] Thus, the use of a mineral oil carrier resulted in a lower
viscosity paste with higher weight % solids than analogous samples
made with a water carrier alone. This type of formulation could be
used as-is by adding it directly to a tooth extraction socket and
by allowing it to hydrate in vivo without textile impregnation and
without pre-masticating with water. This would result in the
highest BUP dosage delivery potential for the formulation.
Alternatively, a formulation like Formula #14B could be optionally
pre-masticated with water, and then deployed in its hydrated state
if so desired. Finally, the lower relative viscosity of Formula
#14B compared to Formula #14A can also render it as useful for
subsequent hemostat textile-impregnation and reinforcement as
demonstrated in Example 6 below.
TABLE-US-00014 TABLE 5-1 Calculations of pertinent composition
information for Formula #14B from the present example and Formula
#14A from Example 4. Although placebo PLGA microspheres were used
in preparing these samples, calculations were performed to estimate
a theoretical dosage of BUP delivery to a tooth extraction socket
having 0.55 cc volume and assuming that the PLGA microspheres were
loaded with 50% BUP by weight. These dosage delivery estimates are
for an embodiment wherein the formulation comprises an oil carrier
that does not cause gelation, and wherein BUP-loaded PLGA
microspheres are suspended in the formulation to yield a compliant
device that can be used for placement into a tooth extraction
socket for subsequent in vivo hydration. Formulation 14A 14B Target
Bupivacaine Dose over 4-Day Period (grams) 0.206 0.177 Estimated
Tooth extraction socket Volume (cm.sup.3) 0.55 0.55 Estimated
density of mixture (g/cc) 1.1 1.1 Estimated grams of "mixture"
delivered to tooth extraction socket (g) 0.605 0.605 Wt. %
Bupivacaine in microspheres 50% 50% Estimated weight of drug-dosed
microspheres in "mixture" (g) 0.41 0.35 Estimated % Total Solids in
"mixture" 83.72% 72.36% Wt. % Microspheres dispersed in "mixture"
67.98% 58.51% Wt. % gelatin in "mixture" 15.74% 13.85% Wt. %
wax/oil in "mixture" portion of formula 16.28% 27.64% Estimated
Total Gelatin + PLGA/BUP Solids in "mixture" Delivered to
0.506535787 0.437807676 Cavity (g) Estimated wt. gelatin delivered
to cavity (g) 0.10 0.08 Estimated wax/oil weight delivered to tooth
extraction socket (g) 0.10 0.17 Estimated wt. % drug-dosed
microspheres in dry formula (excluding 81.20% 80.86% wax/oil)
Estimated wt. % gelatin in dry mixture (excluding wax/oil) 18.80%
19.14% Ratio of oil to gelatin 1.03 1.99 Gelatin Type GLBG GLBG Oil
Type mineral oil mineral oil Sphere distribution type 70/30 w/w
70/30 w/w 42.7/3.4 micron 42.7/3.4 micron
Example 6. Preparation of a System Using a Cellulose Textile
Impregnated with a Formulation Comprising GLBG, PLGA, and Mineral
Oil
[0251] This example describes the composition and preparation of a
fiber reinforced composite device comprising a cellulose hemostat
textile as the fibrous component and a compliant formulation as the
carrier for active ingredients, wherein the formulation comprises
an oil, gelatin binder, and PLGA microspheres, and wherein the
formulation is impregnated into the interstitial spaces of the
fibrous textile. The cellulose fiber textile in this example was a
commercially available hemostat known as SafeGauze.RTM.
Hemostat.TM. Topical Hemostatic Dressing (AMD Medicom, Inc.). The
SafeGauze cellulosic textile was observed to be a loosely woven
mesh-like material, and it was also observed to have ample
interstitial space for impregnation and filling with a compliant
formulation like Formula #14B as described in Example 5.
[0252] In the first step, a single layer of SafeGauze textile was
weighed into a tared beaker at 0.1240 g. The textile was determined
to have unfolded rectangular dimensions of approximately 3.7
cm.times.1.7 cm. Next, 0.6 g of Formula #14B was added to one side
of the textile to prepare a fiber-reinforced composite. Note that
0.6 g of the mixture is estimated to deliver about 177 mg
bupivacaine to the tooth extraction socket as shown in Table 5-1.
The mixture was spread with a spatula to form a bilayer comprising
rectangular fibrous textile on one side and Formula #14B on the
other. The long leg of the rectangular textile was then folded over
and onto the Formula #14B mixture, and the assembly was gently
kneaded to insure filling of the interstitial spaces of the textile
on both sides of the fold. The resulting structure was nearly
square (approximately 1.8 cm.times.1.7 cm), comprising an
impregnated textile folded over and onto itself with both sides
being cohesively held together by the Formula #14B impregnated
therein.
[0253] In a separate demonstration, a similar bi-layer assembly was
prepared, but this time a second textile layer was placed on top to
create a tri-layer comprising an interlayer of the Formula #14B
paste surrounded by two fibrous textile outer layers. The
rectangular tri-layer was then folded 3 times over to successfully
compress the Formula #14B paste into the interstitial spaces of the
two textiles.
[0254] In another demonstration of the concept, the 1.7
cm.times.1.8 cm impregnated fiber textile that was prepared with a
single SafeGauze textile was folded again a second time upon
itself. For comparison, two neat SafeGauze textiles were held
together and were then folded twice over one another. This type of
folding procedure with two neat textiles was similar to that which
would be used by a clinician in preparing the SafeGuaze hemostat
for deployment into a tooth extraction socket. Importantly, each of
the folded structures qualitatively appeared to occupy similar
volumes. Thus, it follows that a single SafeGauze textile,
impregnated with 0.6 g of a formulation to form a composite
reinforced controlled release delivery system, could be readily
deployed to fill a tooth extraction socket. In addition, it was
also qualitatively observed that a single layer of the woven
textile was more than sufficient to reinforce the Formula #14B
formulation. This shows that it is possible to achieve
composite-like reinforcement and to maintain mechanical integrity
while simultaneously allowing for minimization of occupied volume.
In addition, it is conceivable that volume could be further
minimized by using lower density woven textiles, or by using random
non-woven fibers if so desired. Moreover, improved kneading and
pressing procedures could be employed to ensure that all of the
non-occupied space within the porous textile becomes completely
occupied by the amalgamized formulation.
Example 7. Hydrating a Cellulose Textile Impregnated with a
Formulation Comprising GLBG, PLGA, and Mineral Oil
[0255] The delivery system from Example 6, comprising a single
textile impregnated with 0.6 g of Formula #14B, was folded over
three times and was kneaded again to insure filling of the
interstitial spaces with the Formula #14B mixture. The resulting
composite was permitted to age for 1 month under ambient
conditions. No changes in relative compliance or compressibility
were observed after this period of aging.
[0256] In a separate test, a sample of the SafeGauze cellulose
textile was observed to be soluble in water, and when it was placed
in contact with water, it was noted to immediately consolidate into
a sticky mass. Importantly however, the SafeGauze textile material
was observed to remain intact within the composite after 1 month of
being impregnated with Formula #14B, thereby indicating that the
delivery system exhibits good shelf-stability.
[0257] In the next step, water was gradually added to hydrate the
impregnated formulation, with the total water addition equating to
a 2 to 1 ratio by weight of water to the GLBG component within the
mixture. When water was initially placed on top of the folded
textile, the delivery system did not wet immediately. However,
after a short period of time, the entire matrix of textile and
Formula #14B was observed to consolidate, and it was easily kneaded
into various shapes. The delivery system exhibited high compliance
and formability. This suggests that the impregnated textile could
be added directly to the tooth extraction socket to hydrate in
place, or it could alternatively be hydrated with water first, and
then placed into the tooth extraction socket.
[0258] These results also show that it should be possible to also
create different types of formulations for textile impregnation,
including for example, formulas with a low melting wax, formulas
with oil/wax blends, or even with lower Tg control-release polymers
(e.g., lower than the Tg of the PLGA. Simple pressing processes can
be used to pre-consolidate the textile with drug-loaded
microspheres under ambient conditions. The ability to process under
ambient or near-ambient conditions is particularly advantageous for
situations where active ingredients are temperature sensitive.
[0259] Optionally, gelatin binder may be omitted from the
formulation to thereby allow the cellulose textile component to
become the binder for the PLGA microspheres when the delivery
system is hydrated. Omission of the gelatin binder would also make
more "room" for higher levels of bupivacaine-loaded PLGA
microspheres, resulting in higher possible BUP delivery dosages in
volume restricted applications.
Example 8. Formulations Comprising Wax, Oil, GLBG, and PLGA for
Impregnating into a Cellulose Textile
[0260] Use of a wax together with the oil can lead to a further way
of modulating and controlling the rheological characteristics of a
delivery system. Identifying a wax-type, determining the optimum
weight ratio of wax to oil, and the optimum level of wax plus oil
for textile-impregnation required consideration of several factors,
including: 1) the compliance characteristics of the resulting
formulation; 2) the cohesive strength of the formulation; 3) the
hydration rate of the formulation upon exposure to fluids in vivo;
3) the time-dependent mechanical property characteristics of the
formulation during the in vivo hydration process; 4) the
conduciveness of the formulation to textile impregnation during
manufacturing (e.g., solvent-free, minimal pressure, minimal
temperature, textile wettability, etc.); 5) the optional capacity
for the formulation to be pre-hydrated with water before insertion
into the tooth extraction socket if so desired; and 6) the capacity
for the formulation to be delivered with or without a fibrous
textile reinforcing component.
[0261] The present example describes the preparation of a
formulation for delivery as a reinforced composite. Part-1i
describes the use of optional waxes as rheology modifiers for an
oil carrier. Part-2 describes the first step in preparing
formulations where the wax-modified oil carriers from Part-1 are
mixed together with gelatin binder to form amalgamated dispersions.
Part-3 describes the step of mixing and dispersing PLGA
microspheres into the dispersions from Part-2 to yield formulations
for either stand-alone deployment via in vivo-hydration, for
deployment after hydration, or for use in forming fibrous
reinforced composite devices. Part-4 describes the use of the
formulation from Part-3 to prepare a reinforced composite delivery
system for subsequent deployment. Finally, Part-5 illustrates the
optional hydration and mastication of the hydrophobic device from
Part-4 with water prior to deployment.
Part-1. Testing Wax/Oil Mixtures (83.33% Mineral Oil+16.67%
Wax)
[0262] Sample 19-1: 5/1 Mineral Oil to Paraffin Wax 1 g of
household paraffin wax (Gulf Wax, distributed by Royal Oak
Enterprises, LLC, Roswell, Ga.) was weighed into an aluminum pan
and then 5 g of Aldrich Heavy Weight Mineral Oil (CAS 8020-83-5)
was added to yield a 5 to 1 ratio by weight of oil to wax. The
mixture was heated on a hot plate for about 10-20 seconds at 175
degrees C. while stirring with a metal spatula until the wax was
melted to yield a clear homogeneous solution. At that point, the
solution was removed from the hot plate and was allowed to set idle
under ambient conditions. Within 10 minutes, the solution became an
opaque heterogenous dispersion of uniformly suspended wax
crystallites. The mixture had the consistency of a soft spreadable
gel.
Sample 19-2: 5/1 Mineral Oil to Beeswax
[0263] 1 g bees wax (Aldrich, CAS 8012-89-3, cat. #243248, yellow,
melt point 61-65 degrees C.) was mixed with 5 g mineral oil. Using
the same procedure as described above, the solution forms a gel,
but with slightly higher viscosity than Sample 19-1.
Sample 19-3: 5/1 mineral oil to carnauba wax no. 1 yellow
[0264] 0.75 g carnauba wax (Aldrich, CAS 8015-86-9, cat. #243213,
yellow, 82-86 degrees C. melt point) was mixed with 3.75 g mineral
oil. Using the same procedure as described above, a gel formed, but
with higher viscosity than both Samples 19-1 and 19-2. The Sample
19-3 solution was the fastest to recrystallize.
[0265] The viscosities of each of the sample gels can be modulated
by changing the ratio of oil to wax. It is also possible to mix the
waxes, or alternatively to mix pre-formed oil/wax gels of each type
at different ratios. One advantage of mixing different wax-types
together is that it can be possible to modulate viscosity with
multiple combinations of waxes, while simultaneously maintaining a
constant ratio of oil to total wax in the resulting gel. In this
way, lower viscosities can be achieved without having to increase
the level of oil. In addition, by mixing different wax types
together, it can also be possible to minimize the percentage of oil
or the percentage of any other type of low molecular weight
component that is used in the mixture.
[0266] The gels of the types prepared in this example can be used
as carrier components for binder materials that are used in
preparing formulations analogous to Formula #14B from Example 5. In
turn, vehicles incorporating gel carriers can be used in preparing
composite reinforced delivery systems like those prepared in
Example 6. In the next steps, select gels from Part-1 of the
present example will be used to prepare formulations. The gels will
be substituted for the equivalent weight of mineral oil that was
used in preparing Formula #14B in Example 5.
Part 2. Mixing the Wax/Oil Cakes/Gels with Powdered Great Lakes
Bovine Gelatin (GLBG)
[0267] Each of the wax/oil mixtures from Part 1 were separately
melt-dispersed with GLBG over a hot plate for about 10 seconds
(T=175 degrees C.) while stirring with a spatula. The dispersions
were removed from heat and allowed to cool and solidify while
continuing to stir under ambient conditions. The recrystallization
rate was fastest for the highest melting point wax. The final mixed
composition was 55.51% by weight MO, 11.10% by weight wax and
33.39% by weight GLBG.
Sample 23-1. 5 g Mineral Oil+1 g Paraffin Wax+3.007 g GLBG
[0268] Using the procedure described above, GLBG was added to the
19-1 gel from Part-1 to form sample 23-1. In mixing sample 19-1
with GLBG, the resulting 23-1 amalgam provides the same effective
weight ratio of oil-phase to gelatin that was used in creating
Formula #14B from Example 5, where 0.2039 g mineral oil was added
to 0.1022 g GLBG. The weight ratio of oil phase (wax+oil) to
gelatin was 1.995.
Sample 23-2. 5 g Mineral Oil+1 g Beeswax+3.007 g Great Lakes Bovine
Gelatin
[0269] Using the procedure described above, GLBG was added to the
19-2 gel from Part-1 to form sample 23-2. In mixing sample 19-2
with GLBG, the resulting 23-2 amalgam provides the same effective
weight ratio of oil-phase to gelatin that was used in creating
Formula #14B from Example 5, where 0.2039 g mineral oil was added
to 0.1022 g BG. The weight ratio of oil phase (wax+oil) to gelatin
was 1.995.
Sample 23-3. (3.75 g MO+0.75 g Carnauba Wax)+2.255 g Great Lakes
Bovine Gelatin.
[0270] Using the procedure described above, GLBG was added to the
19-3 gel from Part-1 to form sample 23-3. In mixing sample 19-3
with bovine gelatin, the resulting 23-3 amalgam provides the same
effective weight ratio of oil-phase to gelatin that was used in
creating Formula #14B from Example 5, where 0.2039 g mineral oil
was added to 0.1022 g BG. The ratio of oil phase (wax+oil) to
gelatin was 1.995.
[0271] Upon cooling, the 23-3 carnauba wax mixture formed a solid
cake, whereas the 23-2 beeswax and 23-1 paraffin mixtures formed
spreadable gels. The 23-2 beeswax mixture was also higher in
viscosity than the 23-1 paraffin mixture.
Part 3. Adding PLGA Microspheres to the Wax/Oil
Cakes/Dispersions
[0272] The amalgamated dispersions from Part-2 were mixed in a
subsequent step with placebo PLGA microspheres to create
formulations analogous to Formula #14B as described in Example 5.
For the case of Formula #14B, the weight ratio of (wax+oil+gelatin)
to PLGA microspheres was 0.573, and 0.6 g of the Formula #14B
vehicle was impregnated into a single SafeGauze textile.
[0273] In view of Formula #14B, 0.3061 g each of samples of 23-2
and 23-3 were pre-weighed into separate 10 ml beakers. In a
separate step, placebo PLGA microsphere mixtures were pre-weighed
into two separate 10 ml beakers, with each containing 0.1327 g of
3.4-micron and 0.2990 g of 42.7-micron PLGA microspheres, a weight
ratio of large to small microspheres of about 70/30. The
pre-weighed microspheres were then added to the 10 ml beakers for
mixing. The total weight of each formula when mixed was 0.7378 g,
comprising 58.51% by weight of PLGA microspheres and 41.49% by
weight of amalgamated dispersion (i.e., the combined wax+oil+GLBG
mixtures from Part-2). Said another way, the composition of each
formulation included 58.51% by weight PLGA microspheres, 4.61% by
weight of wax, 23.03% by weight of mineral oil, and 13.85% by
weight of GLBG.
Formula #14C
[0274] Sample 23-2, a soft gelatinous dispersion containing beeswax
with MO and GLBG, was spatula-stirred and was weighed into a 10 ml
beaker. The pre-weighed PLGA powder was added and the mixture was
kneaded with a spatula. The resulting mix was a surprisingly tacky
& soft, dough-like material, which indicates that the
percentage of PLGA and hence the potential BUP dosage level could
be increased if so desired. Based on the compliance characteristics
of Formula #14C, the formulation could be directly deployed as a
drug delivery device to hydrate in vivo, or it could be optionally
hydrated for subsequent deployment. Also, based on the relative
compliance characteristics of Formula #14C, the level of total wax
plus oil could optionally be reduced to allow for an increase in
PLGA and BUP dosage levels, or the ratio of mineral oil to beeswax
could be reduced if so desired.
Formula #14D
[0275] Sample 23-3, a soft cake-like dispersion containing carnauba
wax with mineral oil and GLBG, was kneaded into a thick paste,
which was qualitatively higher in viscosity than its Sample 23-2
beeswax counterpart. The pre-weighed PLGA powders were added to
sample 23-3. More mechanical energy was required to knead the
material, and the resultant mixture had more wax-like consistency
than Sample 14-C. It was qualitatively higher in relative
viscosity, yet it displayed good compressibility. Like its Sample
14C counterpart, the Sample 14D formulation could be optionally
deployed for in vivo hydration, or it could be optionally hydrated
for subsequent deployment.
Part-4. Impregnating Cellulose Textile with Vehicles from
Part-3
[0276] There was no qualitative change in the viscosity or hardness
of the Samples 14C and 14D after sitting for 24 hours under ambient
conditions. Cellulose textiles were separately weighed for each
formula mixture (SafeGauze, weight=0.1251 g). Next, 0.625 g of each
mixture was separately added to 1/2 the area of each textile's
rectangular surface. Each textile was then folded over its
respective vehicle mixture, and the composites were gently kneaded
by hand to achieve textile impregnation. Squeeze-out material was
removed by cutting with a spatula. Each impregnated textile was
then re-opened, and additional vehicle was added for the purpose of
exceeding the target-weight of 0.605 g. The total weight of each
vehicle at this point was 0.621 g. Each textile was folded over
again, and then was gently kneaded for a second time. The excess
squeeze-out was cut away with a spatula until the 0.605 g vehicle
target weight was achieved. The composites were then stored under
ambient conditions for future hydration.
Part-5. Hydration of the Impregnated Cellulose Textiles from
Part-4
[0277] After a little more than one month of storage under ambient
conditions, the textiles that were impregnated with the
formulations of Samples 14C and 14D were observed to still be
flexible and compliant, and they had remained qualitatively
unchanged. The impregnated formulas contained 13.85% by weight of
GLBG, which equates to 0.0831 g of GLBG per 0.6 g, where 0.6 g
represents the approximate weight of a formula that is impregnated
into each of the textiles. Note that the net weight of the
impregnated textiles was approximately 0.75 g. In keeping with the
addition of water at a 2/1 weight ratio of water to gelatin that
was used in the prior hydration of the Formula #14B-impregnated
textile in Example 7, 0.1662 g of water was separately added with a
syringe to small weighing boats with each boat containing one of
the impregnated textiles. The formulations were masticated by hand
to yield tacky, highly compliant, dough-like materials. The
effective water to device weight ratio was only about 0.2 parts
water per unit weight of the device. Neither of these devices
exhibited oil-phase exudation as the hydrophobic phase remained
emulsified and stable within each of the hydrated devices.
[0278] In a second part of the experiment, 0.083 g of additional
water was added to the already hydrated Formula #14D impregnated
delivery system, bringing the water to gelatin weight ratio to
approximately 3/1 (w/w), and the effective water to device weight
ratio to approximately 0.33/1 (w/w). The composite was masticated
by hand, and the water was successfully entrapped within the
composite with no evidence of oil or water phase separation. During
mastication, the mixture was tacky and highly compliant, and its
compliance characteristics were qualitatively analogous to those of
neat Surgifoam when Surgifoam is mixed with 3 parts water by weight
to 1 part Surgifoam by weight.
[0279] Thus, formulations compromising hydrophobic components can
be made to have tactile characteristics that are equivalent to
those of other commercially acceptable devices. Moreover,
equivalent characteristics can be achieved with significantly less
water per unit weight of device. Aside from having the benefit of
being usable with less volume-occupying water in an already
volume-restricted application, this water-absorbing feature also
offers the opportunity for controlled dilution of the formulation,
if so desired. For example, if the formulation is manufactured with
an upper-limit dosage of active bupivacaine ingredients, it can
then be diluted to reduce dosages to the degree necessary for the
patient, simply by adding more water to a single type of
manufactured unit. Aside from the manufacturing advantages, such as
minimizing product types and inventory by manufacturing a single
type of formulation, the clinician can simply control dosages by
having the choice of either employing maximum dosage via in vivo
hydration of the delivery system within the tooth extraction
socket, or by diluting the dose via addition to the formulation of
volume-occupying water to a prescribed level, followed by
masticating and cutting to the necessary weight for reaching the
prescribed dosage target, while simultaneously maintaining an
adequate volume-fill factor.
[0280] In yet another step, the hydrated Formula #14D impregnated
device (3/1 sample) was mixed with yet an additional 0.0831 g
water, bringing the water to gelatin weight ratio to 4/1 (w/w), and
the effective water to delivery system weight ratio to
approximately 0.44/1. The formulation was again masticated by hand,
and the water was successfully entrapped in the formulation with no
evidence of oil or water phase separation. During mastication, the
mixture remained tacky and highly compliant.
[0281] Thus, formulations comprising hydrophobic components have
the surprising capacity to absorb hydrophilic fluids, like water,
without undergoing macroscopic phase separation. Moreover, unlike
the formulations comprising hydrophilic components as described in
Example 2, the formulations shown in this example are
erosion-resistant when they are submerged in water under static
conditions, as demonstrated in Examples 11 and 12 below.
[0282] The serendipitous discovery of formulations that resist
erosion while simultaneously allowing for water absorption is both
fortuitous and desirable. This dual capability is what facilitates
in vivo hydration of the formulation on the one hand, while
simultaneously limiting erosion and macroscopic deterioration on
the other. The hydration of the formulation allows for the
diffusive ingress of water with the simultaneous diffusive egress
of molecular-level ingredients, like BUP, to the surrounding
tissues. The fact that this happens without macroscopic phase
separation and without appreciable erosive-deterioration of the
formulation itself is not only unexpected, it is desirable and
beneficial from the standpoint that the viability of the device
relies on its ability to maintain its long-term in vivo cohesive
integrity, and on its ability to simultaneously facilitate the
sustained release of active ingredients. This surprising dual
capability for water-absorption and static erosion resistance is
demonstrated in Examples 12 and 13 below using in-vitro water-soak
experiments together with UV spectroscopy.
Example 9. Drug Delivery Devices Comprising Cellulose Materials
Impregnated with Formulations Compromising Hydrophobic Components
for In Vivo Applications
Part-1. Preparation of the Formulation for a Drug Delivery
Device
[0283] Using the procedures outlined in Example 8, a version of the
Formula #14C was prepared for textile impregnation studies, and for
evaluation during an in vivo porcine study to test the physical and
handling efficacy of the delivery device. In this example, the
Formula #14C was re-designated as Formula 14C-2 owing to the use of
different lots of PLGA particles separately prepared by SWRI and
use of light weight mineral oil in place of heavy weight mineral
oil.
Materials.
[0284] The materials used for the hydrophobic formula preparation
included the following:
1. Mineral oil (MO), white, light, Aldrich Chemical, cat.
#33,077-9, CAS 8042-47-5; 2. Beeswax (BW), Aldrich, CAS 8012-89-3,
cat. #243248, yellow, melt point 61-65.degree. C.; 3. Bovine
gelatin (GLBG) powder, Great Lakes Gelatin Company, Grayslake,
Ill., type B (bovine), unflavored Kosher beef hide, 88-92% protein,
US Pharmacopeia consumer grade; 4. Poly(lactic-co-glycolic acid)
microspheres (PLGA), Southwest Research Institute (SWRI), Resomer
RG504 (Evonik 50/50 grade), spray dried from a solvent solution,
sample designation NB:18-0202-015-15 & -16, particle size
(D50)=5-micron, surface area 1.36 m.sup.2/g; 5.
Poly(lactic-co-glycolic acid) microspheres (PLGA), Southwest
Research Institute (SWRI), Resomer RG504 (Evonik 50/50 grade),
dissolved with solvent into solution, emulsified in a carrier,
solvent extracted and dried, sample designation NB:18-0202-015-14
& -17, particle size (D50)=41-micron, surface area 0.153
m.sup.2/g.
Formula 14C-2 Formulation Preparation Procedure and
Composition.
Step 1.
[0285] The 19-2 MO/BW premix (83.33% by weight mineral oil+16.67%
by weight beeswax) as described in Example 8 was prepared. Solid
beeswax and liquid mineral oil were weighed and placed together
inside of tared aluminum weighing pans. The mixture was heated over
a hot plate having a surface temperature of 175.degree. C. while
stirring with a metal spatula until the wax was melted to yield a
yellowish homogeneous solution. Mix time was about 30 to 45 seconds
until the wax was melted. At that point, the solution was removed
from the hot plate and was allowed to set idle under ambient
conditions. Within 10 minutes, the solution became an opaque
heterogenous dispersion of uniformly suspended wax
micro-crystallites. The mixture had the consistency of a soft
spreadable gel.
Step 2.
[0286] The 23-2 MO/BW/GLBG suspension (55.51% by weight MO+11.10%
by weight wax+33.39% by weight GLBG), referred to herein as the
binder phase, was prepared using procedures similar to those as
described in Example 8. GLBG powder was separately weighed and was
then spatula-stirred under ambient conditions into an aliquot of
the 19-2 gel from step 1. The suspension was then heated over a hot
plate in an aluminum pan while spatula stirring for approximately
30 seconds using a hot plate surface temperature of 175.degree. C.
until the micro-crystallites of the gel were melted. When the gel
phase of the suspension was melted, the pan was removed from the
hot plate, and the dispersion was continuously spatula-stirred
under ambient conditions until the gel phase (oil+wax)
recrystallized to yield a homogeneous suspension of GLBG powder
within a continuous gel matrix phase. This became the binder phase
vehicle for subsequent dispersion of PLGA microspheres, which were
non-drug placebo types in this example.
Step 3.
[0287] The Formula 14C-2 formulation was prepared by dispersing
placebo PLGA microspheres into the 23-2 gel matrix phase from step
2 using procedures similar to those outlined in Example 8. The
target ratio of the two PLGA particle size distributions was
approximately 70/30 w/w 41-micron and 5-micron particles. In the
first step, the 5-micron particles were weighed into a tared
plastic beaker, to which the requisite weight of the 41-micron
particles was then added. The two particle size distributions were
dry-mixed using a spatula. In the next step, the 23-2 gel matrix
phase dispersion from step 2 above was added to the beaker
containing the dry PLGA particles, and the mixture was masticated
using a spatula until a homogeneous dispersion was created. The
final composition contained 41.49% by weight binder (i.e., mineral
oil, beeswax, bovine gelatin) and 58.51% by weight of the PLGA
microspheres. The entirety of the Formula 14C-2 formulation on a
weight % basis is provided in Table 9-1.
TABLE-US-00015 TABLE 9-1 Final 14C-2 Mixture Composition Wt. %
Mineral Oil 23.03% Beeswax 4.61% Bovine Gelatin 13.85% 5 um PLGA
microspheres 17.99% 41 um PLGA microspheres 40.53% TOTAL
100.00%
Part-2. Preparation of Impregnated Fiber-Reinforced Composites for
Use as Drug Delivery Devices.
[0288] Formulations comprising hydrophobic components, such as the
formulation embodied in Formula 14C-2 in Example 9 or others as
described in Examples 3 through 8, can be formulated for in vivo
use in at least four different ways, including for example:
option-1) as a stand-alone device without fibrous reinforcement,
where the formulation is masticated with water-based fluids, such
as saline solution, plasma, etc., before insertion into a tooth
extraction socket; option-2) as a stand-alone device without
fibrous reinforcement, where the device is not masticated with a
fluid, but instead is allowed to hydrate in vivo via static
diffusion processes after being placed within the tooth extraction
socket; option-3) as a device wherein the formulation is first
reinforced with fibrous material, such as knitted, woven or
non-woven cellulose fibers or random cellulose fibers, and then is
masticated with water-based fluids before insertion into a tooth
extraction socket; and option-4) as a device wherein the
formulation is reinforced with fibrous material and is not
masticated with a fluid, but instead is allowed to hydrate in vivo
via static diffusion processes after being placed within the tooth
extraction socket.
[0289] Any one of these four options could be used for drug
delivery. Option-4 is of particular interest for several reasons
pertaining to end use convenience and efficacy. For example, there
is a desire among clinicians to have a device that minimizes the
need for time-consuming processes such as mastication, or other
forms of special handling for deployment. In such an instance, it
would be necessary for the device to exhibit sufficient compliance
for moldability without having to mix with water-based fluids,
while simultaneously having the ability to retain its cohesive
properties during handling, during deployment, and during end use
after deployment.
[0290] A stand-alone option-2 version could be formulated to also
achieve this objective since the formulations comprising
hydrophobic components can be made sufficiently compliant without
the need for mastication with fluids. However, there are several
added advantages of using option-4 that not only help to meet the
handleability needs of clinicians, but also provide synergistic
performance characteristics that satisfy other clinical needs. For
example, by reinforcing the formulation with fibrous material, a
composite is created wherein the formulation is mechanically
reinforced, thus facilitating the optional use of a formulation
that is formulated with less binder phase and with more PLGA
particles than would otherwise be possible without fiber
reinforcement. This helps to satisfy the need for higher drug
dosage deployment when so desired, without experiencing the
deleterious effects on cohesive strength that would otherwise
accompany any diminution in the percentage of binder. The reduction
in the binder results in a decrease in cohesive strength, which can
be more than compensated for by the use of fiber reinforcement.
Fiber reinforcement can also facilitate the use of higher oil
levels in the formulation. Hence the use of lower viscosity
formulas for ease of manufacturing and for ease of deployment in
vivo can be achieved without experiencing the deleterious effects
on cohesive strength that would otherwise accompany any reduction
in the higher molecular weight components of the binder phase.
[0291] With these types of fiber-related factors in mind, four
distinctly different cellulose-based hemostats were chosen for
comparative use in this example. They were chosen so as to not only
demonstrate the flexibility in choice of applicable materials, but
to also demonstrate the importance of the impact of the fiber
member on handling and efficacy during end use as demonstrated
during an in vivo porcine study as described in part-3 of Example
9.
[0292] The fibrous products that were used in this example are
described in Table 9-2. The sample of SafeGauze was in the form of
a rectangular textile and served as a geometric template for
fashioning the other comparative fibrous materials. Each of the
comparative fibrous materials were purposely pre-cut to have
rectangular dimensions similar to those of the SafeGauze product,
and then the samples were weighed to determine the relative
differences in bulk density among the product types. These data are
provided in Table 9-3.
[0293] Aside from the differences in bulk densities and stiffness,
the commercial hemostats were also chosen for their representative
differences in wetting and solubility characteristics. For example,
the SafeGauze product is known to dissolve into a gelatinous
material when it encounters water or body fluids. By contrast, the
Surgicel Original and Nu-Knit products are known to react and
transform much more slowly than SafeGauze. This type of difference
in solubility, wetting, and diffusion characteristics is known by
those skilled in the art to be a function of several chemical and
physical factors, including for example, the degree of oxidation of
the cellulose material, the molecular weight distribution of the
cellulose material, the fiber bundle densities, the knit densities,
and the total fiber surface area per unit volume.
[0294] These types of differences can be important for the end use
in that the mechanical properties and adhesion characteristics can
be influenced both during initial deployment of the device, and
during protracted use under static conditions in vivo. For example,
a more water-soluble fiber might facilitate faster initial wetting
of the tissues within the socket cavity, but if the fibrous
structure dissolves too quickly, the composite's mechanical
properties, such as erosion resistance, might change too quickly as
a function of time under static conditions. Conversely, a less
soluble fibrous member might help the composite to retain its
mechanical characteristics for longer periods of time under static
conditions, but possibly at the expense of less than optimal
handling characteristics during initial deployment of the device.
For example, if the fibrous material is too stiff, owing to a high
knit density or to a slow reaction with fluids, the initial
handling characteristics and initial cavity wetting characteristics
can be less than optimal. If the fibrous member is too slow to
react with body fluids, initial adhesion characteristics might also
be less than optimal.
[0295] As one aspect of this invention, it can be appreciated that
the choice of the fibrous member for the composite delivery device
is an important one, and that the material can be tuned to the
application by controlling the degree of oxidation which affects
solubility, by controlling the molecular weight of the cellulose,
by controlling the fiber surface area per unit volume, by
controlling the fiber bundle density, by controlling the bulk knit
density, etc. Aside from these tunable factors, it is also possible
to use a mixture of fibrous member types. For example, the fibrous
composite could be comprised of both a relatively fast-dissolving
type of fiber member, such as SafeGauze, and a relatively
slow-dissolving member, such as Surgicel Original. Use of multiple
fiber types can impart combinations of desirable characteristics,
including faster initial wetting and better initial adhesion during
deployment from the more soluble fiber member, and longer term
composite integrity from the less soluble fiber member during the
in vivo use period associated with dynamic changes in properties
owing to inter-diffusion of tooth extraction socket fluids with the
device.
[0296] With these concepts in mind, the relative differences among
the commercial hemostats as qualitatively listed in Table 9-4 were
strategically used to conceive of and to create 10 sets of
composites for qualitative evaluation during in vivo experiments.
The resulting devices and their qualitative characteristics are
described in Table 9-5.
TABLE-US-00016 TABLE 9-2 Comparative commercial cellulose fiber
hemostats. Commercial Hemostat Tradename Source Form SafeGauze
.RTM. Hemostat .TM. AMD Medicom, Inc. Woven fibrous cellulosic
textile comprised Topical Hemostatic from yarns of carboxymethyl
cellulose Dressing sodium fibers; measured dimensions ca. 1.8
.times. 3.8 cm. SURGICEL .RTM. Original ETHICON .RTM., division of
Low knit density knitted fibrous cellulosic Absorbable Topical
Johnson and Johnson textile comprised from oxidized regenerated
Hemostat cellulose yarns. SURGICEL .RTM. NU-KNIT .RTM. ETHICON
.RTM., division of High knit density knitted fibrous cellulosic
Absorbable Hemostat Johnson and Johnson textile comprised from
oxidized regenerated cellulose yarns. SURGICEL .RTM. FIBRILLAR .TM.
ETHICON .RTM., division of Layered structure of lightweight random
Absorbable Hemostat Johnson and Johnson fibrous bundles comprised
from oxidized regenerated cellulose fibers. SURGICEL .RTM. SNoW
.TM. ETHICON .RTM., division of Structured non-woven fabric, needle
Absorbable Hemostat Johnson and Johnson punched with interlocking
fibers comprised from oxidized regenerated cellulose fibers.
TABLE-US-00017 TABLE 9-3 Measured weights of fibrous products that
were first pre-cut to dimensions similar to those of the
as-received SafeGauze textiles (approximately 1.8 cm .times. 3.8
cm). These weights are relative indications of the bulk densities
of the materials. Note that the relative densities of these
reinforcing components scale with the mass of fiber per topical
square centimeter, which can be calculated by dividing the average
weight by 6.84 cm.sup.2. Average Number Mass fiber Weight of
samples Standard per topical Sample (g) measured Deviation cm.sup.2
Notes SafeGauz 0.115 7 0.011 0.0168 As received, woven textile as
received Surgicel 0.047 10 0.001 0.00687 Knitted textile, pre-cut
to the x-y Original dimensions of as-received SafeGauze. Nu-Knit
0.124 8 0.002 0.0181 Knitted textile, pre-cut to the x-y dimensions
of as-received SafeGauze. Fibrillar 0.115 9 0.009 0.0168 Random
non-woven fiber pack, pre-cut to the x-y dimensions of as-received
SafeGauze, and then cut approximately in half along the z-axis
(thickness) to provide a bulk weight similar to that of
SafeGauze.
TABLE-US-00018 TABLE 9-4 Summary of relative differences among the
fiber types after pre-cutting to the same x-y dimensions of the
as-received SafeGauze product. Sample Qualitative differences
SafeGauze as- Used as the qualitative standard in these
comparisons; exhibits relatively high received (SG) solubility and
gel formation almost immediately upon contact with water (within 5
minutes). Surgicel Original Knit structure slightly more open than
woven SafeGauze; not as stiff as (SO) SafeGauze; approximately 1/3
the bulk density of SafeGauze; significantly less water sensitive
than SafeGauze upon initial contact with water. Exhibits slight
shrinkage within 5 minutes but does not dissolve. Remains intact
for 24 hours when coated with drops of water. Nu-Knit (NK)
Significantly tighter knit structure than Surgicel Original despite
its similar bulk density, and higher in stiffness than both
SafeGauze and Surgicel Original. The tighter knit structure at
similar bulk density to Surgical Original is an indicator of a
difference in fiber bundle structure and/or in net surface area per
unit volume of sample. The water sensitivity is similar to Surgicel
Original (less sensitive than SafeGauze). Fibrillar (FIB) Non-woven
random fiber pack; significantly less contiguous interstitial voids
than either of the other woven structures; more resistant to water
than SafeGauze (i.e., less susceptible to initial water diffusion
than SG), and somewhat more water sensitive than SO and NK.
TABLE-US-00019 TABLE 9-5 Summary of devices that were made,
preparation methods, and their qualitative characteristics both
during and after their preparation. Except where noted otherwise,
textiles were impregnated with the 14C-2 hydrophobic formulation
using procedures similar to those outlined in Example 8. Each
device had final x-y dimensions of approximately 1.8 .times. 1.9
cm. Sample Set, Device Designation, weights of device members
Description of device construction Set-1 SafeGauze rectangular
textile (ca. 1.8 .times. 3.8 cm) 1A textile = 0.1214 g.; 14C-2 =
0.6472 folded in half over approximately 0.60 to 0.65 g 1B textile
= 0.1209 g; 14C-2 = 0.6248 g 14C-2; final x-y dimensions =
approximately 1.8 1C textile = 0.1032 g; 14C-2 = 0.6187 g cm
.times. 1.9 cm 1D textile = 0.1128 g; 14C-2 = 0.6409 g 1E textile =
0.1192 g; 14C-2 = 0.6499 g Set-2 Surgicel Original (SO) textile
(when cut to same 2A - textile wt. = 0.4093; 14C-2 = 0.6495 g; 2nd
dimensions as SafeGauze and when folded in half textile wt. =
0.0488 g over approximately 0.60 to 0.65 g 14C-2) does 2B - textile
wt. = 0.0474 g; 14C-2 = 0.6408 g; 2nd not have the same
interstitial-space capacity to textile wt. = 0.0487 g absorb the
14C-2 formula as SafeGauze, nor does 2C - textile wt. = 0.0437 g;
14C-2 = 0.6208 g; 2nd it have the same mechanical integrity (the
textile textile wt. = 0.0502 g is 1/3 the weight with a slightly
more open knit 2D - textile wt. = 0.0447 g; 14C-2 = 0.6147 g; 2nd
structure). For this reason, a second textile was textile wt. =
0.0537 g used in the construction of this set. A pre-cut SO 2E -
textile wt. = 0.0443 g; 14C-2 = 0.6603 g; 2nd textile was coated,
folded on itself, and finger- textile = 0.0485 g pressed to get
interstitial space impregnation. A second SO textile was then
folded over the first folded component members of the construction
in the cross orthogonal direction, and the composite was
finger-pressed to achieve formula impregnation of the outer SO
textile member B, which had then encapsulated the first folded
member A. The final folded construction (along the z-axis) =
[impregnated SO textile layer B orthogonally positioned to
A]/[impregnated SO textile layer A]/[impregnated SO textile layer
A]/[impregnated textile layer B orthogonally positioned to A]. Note
that the use of two SO textiles still results in a composite with a
lower weight percent of fiber than that of set-1 which was made
with SG. Despite this difference, the set-2 composites were
qualitatively similar in stiffness to the set-1 composites. Set-3
Completely analogous to set 1, but with NuKnit 3A - textile wt. =
0.1312 g; 14C-2 = 0.6115 g (NK) textile (cut to same dimensions as
SG) folded 3B - textile wt. = 0.1307 g; 14 C-2 = 0.6551 g in half
and over approximately 0.60 to 0.65 g 3C - textile wt. = 0.1325 g;
14C-2 = 0.06415 g 14C-2. Note that the rough side of the NK textile
3D - textile wt. = 0.1331 g; 14C-2 = 0.6297 g was coated before it
was folded and impregnated 3E - textile wt. = 0.1328 g; 14C-2 =
0.6336 g by finger-pressing. This composite was qualitatively
higher in stiffness than the comparable composite made with SG
(set-1). Set-4 Fibrillar random fiber patch was cut to the same 4A
- fiber wt. = 0.1037 g; 14C-2 = 0.5991 g x-y dimensions as
SafeGauze textile, and it was 4B - fiber wt. = 0.1187 g; 14C-2 =
0.6458 g then cut in the near-center of the z-axis to 4C - fiber
wt. = 0.1168 g; 14C-2 = 0.5990 g achieve similar weight. The
rectangular slab was 4D - textile wt. = 0.1208 g; 14C-2 = 0.6772 g
folded over onto itself and pressed by hand to 4E - textile wt. =
0.1201 g; 14C-2 = 0.6038 g impregnate the higher surface area
random fibers. This resulted in a less homogeneous macro-structure
when compared to the other sets, where the interior of the
composite sandwich was higher in 14C-2 concentration, and the
exterior of the composite was higher in dry fiber concentration.
Thus, although the bulk weight of set-4 was similar to set-1 (also
with similar wt. percentages of the device members), set-4 was
qualitatively, less malleable, higher in stiffness, and less tacky
than set-1. Set-5 Set-5 was a composite of 14C-2 and Fibrillar 5A -
0.78 g cellulose fibers that were homogeneously 5B = 0.8077 g
blended within the 14C-2 hydrophobic formula 5C = 0.8304 g (ca.
97/3 w/w 14C-2/fiber). In preparing set-5, 5D = 0.8167 g the first
experiment involved taking 0.7067 g 5E = 0.8593 g 14C-2 + 0.0234 of
pre-torn fiber; and masticating it in a plastic weighing boat with
a spatula to final wt. = 0.6956 g. The ratio of Fibrillar/14C-2 =
0.0331; and the compliance of this random composite sample was
slightly higher than the SafeGauze 1A sample. Thus, a decision was
made to use slightly more Fibrillar to achieve higher modulus. In
order to accomplish this, 3.215 g of 14C-2 was initially placed
into a 15 ml HDPE beaker, and was spatula-masticated with 0.15 g
pre-torn Fibrillar fibers (Fibrillar/14C-2 = 0.0466). Mastication
of this larger quantity led to a drier blend, so more 14C-2 was
added (.5549 g) to bring the ratio of Fibrillar to 14C-2 = 0.03978.
With continued mastication, it was still somewhat dry, so an
additional 0.971 g of 14C-2 was back-added (4.7409 total), bringing
the Fibrillar/14C-2 ratio = 0.0316. The process of stirring with
mastication continued to tear the fiber bundles and to produce
enough shear to increase fiber surface area, which further
increased viscosity. However, as opposed to adding more 14C-2, the
composite was cut into approximately 0.8 g aliquots. Each aliquot
was comprised of approximately 3% fiber, and 87% 14C-2. Thus, a 0.7
to 0.8 g aliquot contained about 0.6 to 0.7 g of 14C-2. Set-6 In
preparing the set-6 composites, the NuKnit 6A - SafeGauze textile
wt. = 0.0906 g; 14C-2 = rectangular samples, originally cut to the
size of 0.6812 g; NuKnit textile = 0.0518 g SafeGauze rectangles,
were purposely cut in half 6B - SafeGauze textile wt. = 0.1180 g;
14C-2 = (1.9 cm .times. 1.8 cm), trimmed, and weighed. 0.6638 g;
NuKnit textile = 0.0536 g SafeGauze rectangular samples were evenly
6C - SafeGauze textile wt. = 0.0902 g; 14C-2 = coated with
approximately 0.6 to 0.7 g 14C-2. 0.6423 g; NuKnit textile = 0.0457
g The trimmed NuKnit textile was placed on top of 6D - SafeGauze
textile wt. = 0.1200 g; 14C-2 = a1/2-section of a fully coated
SafeGauze textile. 0.6528 g; NuKnit textile = 0.0511 g The other
half of the coated SafeGauze textile 6E - SafeGauze textile wt. =
0.1049 g; 14C-2 = was folded over and on top of the NuKnit textile.
0.6299 g; NuKnit textile = 0.0545 g The sample was compressed
lightly by hand to assist in impregnating the members. The
resulting construction as dissected through the z- axis = partially
impregnated SafeGauze/14C- 2/partially impregnated
NuKnit/14C-2/partially impregnated SafeGauze. In spite of purposely
reducing the relative weight of the NuKnit member, the resulting
construction was qualitatively stiffer than sets 1 and 3. This was
in part due to the relatively high surface area of the NuKnit
member, which resulted in more 14C-2 absorbance by the NuKnit
center member than the SafeGauze outer-layer members. Consequently,
the z-axis distribution of 14C-2 was more heterogeneous than that
of sets 1, 2, and 3. The purpose of this multi-membered composite
(like that of set-7) was to provide an outer layer of
water-sensitive cellulose for the purpose of imparting fast tissue
wetting and tissue adhesion during deployment. The purpose of the
less water-sensitive inner member was to provide the device with
protracted reinforcement for improved cohesive strength throughout
the duration of its static dwelling within the tooth extraction
socket cavity. Set-7 Surgicel Original (SO) rectangular samples, 7A
- SafeGauze textile wt. = 0.1039 g; 14C-2 wt. = originally cut to
the size of SafeGauze rectangles, 0.6568 g; SO textile wt. = 0.0253
g were cut in half (squares), and weighed. 7B - SafeGauze textile
wt. = 0.1068 g; 14C-2 wt. = SafeGauze rectangular samples were
evenly 0.6523 g; SO textile wt. = 0.0267 g coated with
approximately 0.6 to 0.7 g 14C-2. 7C - SafeGauze textile wt. =
0.1285 g; 14C-2 wt. = The SO textile was placed on top of a 1/2
section 0.6525 g; SO textile wt. = 0.0268 g of coated SafeGauze
textile. The other half of the 7D - SafeGauze textile wt. = 0.1200
g; 14C-2 wt. = coated SafeGauze textile was folded over and on
0.6625 g; SO textile wt. = 0.0265 g top of the SO textile. The
sample was compressed lightly by hand to assist in impregnating the
members. The resulting construction as dissected through the
z-axis: partially impregnated SafeGauze/14C-2/partially impregnated
SO/14C-2/partially impregnated SafeGauze. The more open knit
structure of the SO resulted in better 14C-2 homogeneity along the
z-axis than that which was achieved in the comparable composite
made with NuKnit (set-6). Consequently, this multi-member fibrous
composite was less stiff than set-6, and only slightly stiffer than
sets 1 and 2. The purpose of this multi-membered composite (like
that of set- 6) was to provide an outer layer of water- sensitive
cellulose for the purpose of imparting fast tissue wetting and
adhesion during deployment. The purpose of the less water-
sensitive inner member was to provide the device with protracted
reinforcement for improved cohesive strength throughout the
duration of its static dwelling within the tooth socket cavity.
Set-8 This construction was a bi-layer with 0.6 g to 0.7 8A -
SafeGauze wt. = 0.0521 g; 14C-2 wt. = 0.6226 g of 14C-2 interlayer
material. Layer-1 was a cut g; NuKnit wt. = 0.0611 g sample of
SafeGauze (1/2 of a SafeGauze 8B - SafeGauze wt. = 0.0505 g; 14C-2
wt. = 0.6032 rectangle, 1.9 cm .times. 1.8 cm), and layer-2 was a
g; NuKnit wt. = 0.0558 g NuKnit layer cut to the same dimensions as
1/2 of 8C - SafeGauze wt. = 0.5099 g; 14C-2 wt. = 0.6208 the
as-received SafeGauze textile. The 14C-2 g; NuKnit wt. = 0.0604 g
interlayer was lightly pressed by hand to 8D - SafeGauze wt. =
0.0517 g; 14C-2 wt. = 0.6678 impregnate the members. During initial
g; NuKnit wt. = 0.0618 g evaluation, this construction was targeted
to be deployed with the SafeGauze side down towards the tooth
extraction socket tissue. The top side (NuKnit) was intended to
fold into itself as the device was deployed into the tooth
extraction socket. The top side of the device with NuKnit was
marked with a black dot. Note that this device could optionally be
deployed in the opposite direction. However, the original intent
was to provide better initial tissue wetting via use of a more
water-sensitive outside member (SG). In this sense, set-8
represents a similar but subtly different manifestation of the
set-6 construction. Set-9 This set was completely analogous to
set-8 with 9A - SafeGauze wt. = 0.0612 g; 14C-2 wt. = one
exception: Surgicel Original was used as 0.06585 g; Surgicel
Original wt. = 0.0455 g layer-2, and instead of cutting it to the
same 9B - SafeGauze wt. = 0.0672 g; 14C-2 wt. = 0.6238 square shape
as 1/2 the as-received SafeGauze g; Surgicel Original wt. = 0.0541
g rectangle, a full rectangular piece of SO was used, 9C -
SafeGauze wt. = 0.0617 g; 14C-2 wt. = 0.6225 and it was folded in
half to give it the square g; Surgicel Original wt. = 0.0507 g
shape of layer-1 (this was done because SO is 9D - SafeGauze wt. =
0.0540 g; 14C-2 wt. = 0.6614 only 1/3 the weight of SafeGauze). The
final g; Surgicel Original wt. = 0.0473 g construction as dissected
along the z-axis: SafeGauze/14C-2/SO. This construction was
intended to be initially evaluated by deploying it with the
SafeGauze side down towards the tooth extraction socket tissue. The
top side (SO) was thereby intended to fold into itself as the
device
was deployed into the tooth extraction socket. The top side of the
composite with SO was marked with a black dot. Note that this
device could optionally be deployed in the opposite direction.
However, the original intent was to provide better initial tissue
wetting via use of a more water-sensitive outside member (SG). In
this sense, set-9 represents a similar but subtly different
manifestation of the set-7 construction. Set-10 This set was
analogous to set-2, which was made 10A- Surgicel Original textile
wt. = 0.0490 g; 14C- with two rectangular members of Surgicel 2 wt.
= 0.6072 g Original. However, in this case (set-10), only one 10B -
Surgicel Original textile wt. = 0.0483 g; 14C- rectangular member
of SO was used instead of 2 wt. = 0.6248 g two. In this sense,
set-10 was the analog of sets 10C - Surgicel Original textile wt. =
0.0540 g; 14C- 1 and 3, each having been prepared with 1 2 wt. =
0.6436 g rectangular textile member of SG and NK, 10D- Surgicel
Original textile wt. = 0.0515 g; 14C- respectively. The SO textile
was cut to the same 2 wt. = 0.6115 g rectangular dimensions as
SafeGauze, and it was folded in half over approximately 0.60 to
0.65 g 14C-2. Given the lower density of SO, the construction was
substantially lower in stiffness than sets 1, 2, and 3 (set-2
having been comprised of two SO rectangular members instead of
one).
Part-3. In Vivo Evaluations of the Impregnated Fibrous Composites
for Use as a Drug Delivery Device.
Description of the Test Environment and Experimental Details for
the In Vivo Porcine Trial.
[0297] The device samples described in Table 9-5 were used for this
study. Qualitative notes and observations are provided in Table
9-6. Importantly, although some of the devices have preferable
attributes that differentiate them from others, most of the device
constructions exhibited acceptable utility for the application, and
many of the qualitative differences were consistent with the
previously noted qualitative differences among the devices' fibrous
members (Tables 9-3 and 9-4) and among the devices themselves
(Table 9-5). These results show that despite the identical usage of
the Formula 14C-2 formulation, the macrostructural differences
associated with the different fiber-types and construction methods
led to large differences in performance characteristics. From
handling and initial deployment perspectives, set-2 made with SO,
set-3 made with NK, and set-7 made with SG and SO mixed textile
types exhibited a good overall balance of acceptable performance
characteristics. Although the random fiber set-4 performed well
after deployment, its initial handling characteristics were found
to not be as good as comparable composites that were made with
knitted or woven fibrous textiles, thereby providing an
illustration of the importance of fiber type. Similarly, although
the handleability of set-1 made with water-sensitive SG fibers was
deemed to be good, its fast dissolution and resulting lower
durometer in the tooth socket made it less desirable under
post-deployment static conditions than comparable constructions
made with SO and NK of sets 2 and 3 made with less water sensitive
fibers.
[0298] Devices that were prepared with less water-sensitive fibers,
sets 2 and 3 with SO and NK, respectively, tended to form more
homogeneous and higher durometer composite structures in vivo than
samples made with more water-sensitive fibers such as set-1 with
SG. The use of relatively fast-dissolving, water-sensitive fibers
resulted in a qualitative deterioration in modulus (durometer) as
the device became inter-mixed with cavity fluids under static
conditions. This loss in fibrous reinforcement was deterred by the
use of the more water-resistant fibers. Thus, even when samples
exhibited similar pre-deployment mechanical characteristics as in
sets 1 and 2, the difference in fiber-type led to an extreme
difference in mechanical behavior during the post-deployment
period. This is yet another example of the engineering latitude
afforded by the present invention, and the importance of making the
correct fiber choice for the end use application.
[0299] Another finding was related to the degree to which the
fiber-type either facilitated or deterred the formation of a
homogeneous in vivo composite under static conditions with fluid
components from within the tooth extraction socket. As shown in
prior examples, formulations like Formula 14C-2 have the unexpected
ability to absorb and emulsify hydrophilic fluids without
exhibiting macro phase separation. However, the in vivo
observational trends showed that this capability was sometimes
deterred by the use of SG fibers and was generally enhanced by the
use of SO and NK fibers. Under post-deployment static conditions,
the physical probing of the in vivo composites revealed that sets
made with SO and NK tended to become more homogeneously infused
with blood after short time periods, even within their relatively
hydrophobic central regions, indicating that the fibers had
facilitated in diffusion-assisted mixing of blood components with
the Formula 14C-2 formulation. By contrast, sets that were made
with SG as a fibrous member were generally observed to be more
heterogeneous during the post-deployment period. Hybrid devices
that were made with two fiber types, such as SG and SO, exhibited
combined behaviors with macroscopically visible regions where blood
had become more homogeneously dispersed than in samples made with
SG alone, but also with regions that were more heterogeneous than
those observed in samples made with SO or NK alone.
[0300] One advantage of diffusion-assisted mixing is that the
resulting in vivo composite becomes more homogeneous, and from a
mechanical property perspective, this can help to dissipate
internal cavity stresses over a larger volume fraction of the
socket, thereby helping to minimize surface stresses that could
disrupt protective scab formation. In this sense, it also becomes
possible for the composite to become an integral component of the
protective scab itself, wherein the radial gradient in composition
between the tissue surface and the center of the cavity becomes
more homogeneous. From a drug elution perspective, this also
creates a more homogeneous chemical environment for 2-way diffusion
processes, such as free-base BUP diffusion from PLGA, water
diffusion into PLGA to cause hydrolysis and molecular weight
diminution, diffusion of proton-carriers (i.e., Bronsted acids)
toward free-base BUP molecules, etc. Thus, a homogeneous composite
environment can have a profound effect on chemical efficacy.
[0301] By contrast, when a more heterogeneous environment is
enabled to persist for longer periods, the diffusion
characteristics and hence the chemical efficacy can be made to vary
quite substantially. For example, under heterogenous conditions,
the free-base form of BUP may be much slower to protonate, a
process which renders it more water soluble, which would have the
effect of slowing the bulk rate of release, and thereby the effect
of reducing the bio-availability of the drug at any given time.
[0302] Mixed environments afforded by use of multiple fiber-types
can lead to mixed effects. For example, the more homogeneous
regions could be conducive to faster diffusion and bioavailability,
whereas the heterogeneous regions might serve to release their
active ingredients more slowly, which in essence would render them
as storage vesicles for longer-term release. Thus, by controlling
the choices of, the ratios of, and the geometric placement of fiber
types, it can become possible to impact the global morphology of
the device, and hence the global time-release profile of active
ingredients. As long as sufficient mechanical integrity can be
established and maintained by means of homogeneous infiltration and
diffusion of body fluids into some regions of the device,
heterogeneous vesicles larger in scale than the micron-sized PLGA
particles can be allowed to persist for the purpose of facilitating
longer-term release. Depending on the morphology of the resulting
composite structure, the heterogeneous vesicles could even be used
to impart mechanical benefits like stress dissipation. For example,
if the device is engineered to allow for the fast in vivo formation
of a homogeneous blood-mixed continuous phase containing a
dispersed heterogeneous blood-free phase, the resulting morphology
would be analogous to that of many impact-modified materials such
as certain polymeric blends (e.g., impact modified polystyrene with
a polybutadiene dispersed phase), which benefit from
stress-dissipation owing to their dispersed components.
[0303] Thus, the choice of fiber type, single types and mixed
types, affords surprisingly extreme flexibility for achieving
different morphologies and hence varying degrees of control over
performance attributes ranging from mechanical properties (e.g.,
cohesive integrity and resistance to in-use stresses and erosion),
to chemical properties (e.g., diffusion rates and time release
profiles), and combinations of the two. This type of
macro-structural flexibility affords the opportunity to tune the
delivery device for various end use needs, and to provide the
efficacy characteristics that are desired not only for oral surgery
applications, but also for other applications as well.
TABLE-US-00020 TABLE 9-6 Summary of clinical observations from the
in-vivo porcine study. Set Placement Placement Device Malleability
Device Stability General # Location Time Prep. Handling Placement
In vivo Comments 1a Rt. Maxilla, 8:19am Quick, Liked the Air/Fluid
T0 - 8:19, Massive #1 - Molar Easy, No handling. displacement -
Breaking down socket, Blending good, too quickly, heavy Sticking to
Bleeding coming bleeder socket wall, from edges, Conforms to
Appears socket nicely saturated with blood. T1 - 8:30, Continues to
break down, Saturated with blood. T2 - 8:37, Saturated with blood,
T3 - 8:49, Semi- Liquid T4 - 9:08, Still in place but mushy T5 -
10:28, becoming displaced 1b Lt. Maxilla, 9:37am Handles Material
gets T0 - 9:37, infuses Re-test of #19 - Pre- very well. infused
with with blood and Set #1 to molar the blood melts evaluate
(rapidly), into the socket. within a Good T1 - 10:09, a little more
hemostasis, mushy, but stays nominal Air/Fluid in place socket
displacement - with irrigation and bleeding good T2 - 10:27,
conditions becoming displaced from socket 2a Rt. Maxilla, 8:29am
Minimal More stable - Air/Fluid T0 - 8:29, more One of Dr. #3 -
Molar effort, no not displacement stable during Neshat's top
blending breaking was very placement than 3 favorites. Low/No down
as good, Set 1, seemed impact to quickly as Seemed to solid after
current Set 1 achieve placement. surgical hemostasis T1 - 8:41,
Still a procedures more rapidly solid mass than Set 1, T2 - 8:49,
Not Appears breaking down saturated much with blood, T3 - 9:07,
Some More solid softening noticed feeling than T4 - 9:21, Set 1
after Remains intact placement T5 - 9:53, Solid and in place after
irrigation, blood found within center of mass - good sign 2b No
tooth 10:18am None. Easy Air/Fluid T0 - 10:18, extracted
displacement T1 - 10:23, Blood created Lt. was good, thoroughly
Maxilla beautiful, incorporated Hemostasis deep within was rapidly
device achieved 2c Replaced 10:25am Great Better device 9a
hemostasis stability (No tooth on a big during and extracted -
bleeder post device Lt. Maxilla) placement within an excessively
bloody socket, as compared to Set 1. 3a Rt. Maxilla, 8:36am Minimal
Handling Easy T0 - 8:36, Good #4 - Molar effort, no was insertion,
placement blending acceptable, Air/Fluid qualities and Low/No Good
displacement rapid hemostasis. impact to consistency was very T1 -
8:41, current good, Quick Remains surgical coagulation contiguous
procedures T2 - 8:50, very solid, very similar to Set #2 T3 - 9:07,
most solid T4 - 9:21, Still intact T5 - 9:55, Still in place after
irrigation, blood seen throughout, good handling 3b 10:19am None
Good T0 - 10:19, needed. incorporation T1 - 10:30, of blood noted.
homogenous the way it reacted with the blood, a little stiffer. 4a
Rt Maxilla 8:47am None. Difficult Air/Fluid T0 - 8:47, Heaviest #6
- Pre- Handling, displacement Performs well in bleeding molar
cannot was socket site, press well acceptable T1 - 8:90, Solid,
(through into the stayed in place nasal) socket T2 - 9:23, Mushy,
but does go more break down in. Post- noted than 7a application T3
- 9:55, Still in performed place after well, irrigation handling T4
- 10:01, Good and handling, blood placement within difficult. 5a No
tooth 9:06am None Very soft, Went in T0 - 9:06, Good extracted
sticks to nicely, placement created gloves, Air/fluid qualities,
and furthest displacement rapid hemostasis. toward was T1 - 9:24,
Soft nose in the acceptable, T2 - 9:56, Still in maxilla but not as
there, disrupted good as some some prior T3 - 10:01, Very
prototypes. soft Hemostasis T4 - 10:05, was Broken down achieved
more than quickly. other prototypes 6a No tooth 9:14am None Sticks
to Good T0 - 9:14, Good extracted gloves air/fluid placement
created displacement qualities, and socket in Went into good
hemostasis maxilla socket well. T1 - 9:24, Very good, not breaking
down, clotting well T2 - 9:56, Good after irrigation T3 - 10:02,
Fragmented, not stable 7a No tooth 9:19am none No tackiness Went in
T0 - 9:19, Good extracted to gloves nicely, placement created
Air/fluid qualities and socket in displacement good hemostasis.
maxilla was good, T1 - 9:25, Large Hemostasis socket, doing a was
achieved good job. quickly. T2 - 9:58, Breaking down, did not look
like much coagulation. 7b Socket 10:04am None Very good Air/fluid
T0 - 10:04, One of reused handling, displacement the best for from
"one of was good, handling prototype the best" Hemostasis and
placement. 6a. was achieved T1 - 10:14, Two quickly. phases are
obvious, no blood seen inside prototype 7c Socket 10:22am None
Great Very good T0 - 10:22, Great reused handling hemostasis
placement and from hemostasis. prototype T1 - 10:26, 10a. Doesn't
have blood incorporated as much (as other prototypes) 8a No tooth
9:43am None Falls apart Air/fluid T0 - 9:43, Falls extracted (while
displacement apart while created handling) was good, forming plug
in toward Hemostasis gloved hands. nose was T1 - 10:10, Top within
the achieved, piece falls out maxilla Prototype became firm after
placement in socket 9a No tooth 9:47am None Nice Air/fluid T0 -
9:47, Good extracted handling, displacement placement and created
folded and was very hemostasis socket in went into good, T1 -
10:11, Very maxilla socket well. Hemostasis mushy Feels good. was
throughout, not achieved, good, very gel like. 10a No tooth 9:50am
Easy Mushy to Air/fluid T0 - 9:50, Heavily extracted start with,
displacement Prototype was bleeding site created really was very
mushy during socket in breaking good, placement, good maxilla down.
Hemostasis hemostasis. was T1 - 10:12, achieved, Outside had
clotting, but no blood was found inside prototype. T2 - 10:21, Not
Good
Example 10. Devices from Example 9 Prepared with Isopropyl
Palmitate and Caprylic Triglyceride in Place of Mineral Oil
[0304] Using the procedures outlined in Example 9, the Formula
14C-2 formulation provided in Table 1 from Example 9 was used as a
guide to prepare two analogous formulations with different oil
substitutions for the mineral oil component. In one case, a
formulation designated as 12019-23-1 was made using isopropyl
palmitate in place of mineral oil (Sigma-Aldrich Cat. # W515604;
lot # MKCB9456; >90% isopropyl palmitate; CAS #142-91-6; 298.5
g/mole; melt point reported as 11 to 13 degrees C.; density=0.852
g/ml at 25 degrees C.). In a second case, a formulation designated
as 12019-23-2 was made using caprylic triglyceride in place of
mineral oil (Croda, Inc.; CAS #65381-09-1; Columbus Circle, Edison,
N.J.; tradename Crodamol GTCC). Again, apart from the type of oil,
the compositions and relative weight percentages of all ingredients
were the same as those used in preparing Formula 14C-2.
[0305] Both alternative oil-types led to homogeneous compositions
with no evidence of macro phase separation or oil exudation. During
mixing, the qualitative compliance characteristics were evaluated
and ranked from high to low. 12019-23-1 with isopropyl palmitate
was kneaded to form a homogeneous dough-like mixture with relative
compliance that was qualitatively higher than that of 12019-23-2
with caprylic triglyceride. The compliance ranking from high to low
was as follows: Formula 14C-2 was more compliant than 2019-23-1
which was more compliant than 12019-23-2. Interestingly, this
result shows that with all other things being equal, the simple
substitution of a different type of oil can have a significant
impact on the mechanical properties of the drug carrier
formulation. In this example, the effect was qualitatively similar
to that caused by substitution of a different type of wax, or by
the use of a different wax to oil ratio. Thus, this example
provides further illustration of the versatility in rheological and
mechanical property characteristics that are possible by means of
controlling not only composition percentages, but by also
controlling the chemical nature of the components. In this example,
three different rheological characteristics were achieved by merely
changing the nature of the oil type.
[0306] The impact of the differences in properties from this simple
substitution become more apparent when consideration is given to
the manufacturability and to the efficacy of the final composite
device when the formula is paired with a fibrous member such as one
employed in Example 9. In order to illustrate this, the three
comparative formulations in this example were each separately
impregnated into pre-cut textiles from the Surgicel Original
cellulosic hemostat that was used in Example 9. The textiles were
cut to the same approximate dimensions as those used in Example 9,
and the three formulations were paired with the textiles using the
same approximate weight of carrier formula that was employed in
Example 9.
[0307] Initially, an attempt was made to create devices like those
designated as set-2 and described in Table 9-5 of Example 9 with
two SO textiles. As previously noted in Example 9, when the higher
compliance Formula 14C-2 was impregnated into a single SO textile,
the resulting composite was relatively low in stiffness (see set-10
from Example 9). Moreover, the excess penetration of the Formula
14C-2 into the interstitial spaces of the SO textile necessitated
the use of a second textile to achieve better in vivo performance
as described in Tables 9-5 and 9-6 of Example 9. By contrast, when
attempts were made to create analogous devices with 12019-23-1 and
12019-23-2, the lower relative compliance of these formulas led to
less interstitial penetration, and to higher qualitative stiffness
characteristics, which thereby negated the need for a second
textile component. In essence, the stiffness of devices comprising
either 12019-23-1 or 12019-23-2 with one textile was qualitatively
similar to the stiffness of devices comprising Formula 14C-2 with
two textiles. One advantage of this versatility relates to the
efficacy of the final composite device. Specifically, when using a
textile with relatively low knit density like SO, a device can be
prepared with a lower volume fraction of the cellulosic hemostat
component, and with a higher volume fraction of the formulated drug
carrier simply by changing the chemical nature of the oil component
in the drug carrier formulation. Similar results would also be
possible by changing other factors either alone or in combination,
including for example, the wax type, the oil to wax weight ratio,
and the volume fraction of dispersed solids such as PLGA, and
gelatin. Thus, given the volume restrictions associated with the
end use application, the ability to control compliance
characteristics with these factors can lead to a reduction in the
volume % of cellulose textile in the device and consequently to a
higher dosages of active ingredients per unit volume if so
desired.
[0308] In summary, with increasing formula compliance and with
lower textile density, excessive formula penetration into the
interstitial spaces and a lower net composite stiffness may
necessitate the use of a second orthogonal textile to achieve
acceptable tactility in terms of stiffness and handleability for
certain end use applications. By contrast, lower compliance
formulas do not produce the same degree of interstitial
impregnation under equivalent pressure, and because they are
inherently stiffer, the need for a second textile can be negated.
In this example, the lower compliance formulas that have been
paired with one SO textile member, analogous to set-10 from Tables
9-5 and 9-6 in Example 9, exhibited qualitatively similar stiffness
characteristics to Formula 14C-2 that was made with two SO textile
members. Again, one of the advantages of using a lower compliance
formula with one low density textile instead of two, particularly
when the textile is as low in density as SO, is that the occupied
fibrous hemostat volume fraction can be reduced without necessarily
compromising the types of tactile characteristics that are
important during clinical end use. This can equate to a higher
volume fraction of the drug vehicle, and to higher possible drug
delivery dosages in certain volume restricted end use applications,
as represented by the tooth extraction socket application.
[0309] On the other hand, if the hemostat character of the delivery
device is of particular functional interest, then the device can be
optionally made with higher relative volume fractions of oxidized
cellulose if so desired. This type of composite device would
necessitate the use of a formulation with higher compliance
characteristics. Indeed, this approach was demonstrated previously
in Example 9 with the higher compliance Formula 14C-2 made with
mineral oil as the liquid carrier. Sample set-2, which was made
with two SO textiles, provided better in vivo performance, better
diffusion-assisted mixing with body fluids under static conditions
and better homogeneity within the oral tooth socket than sample
set-10 which was made with only one SO textile.
Example 11--Water Soak Experiment Involving Samples from Example
10
[0310] As mentioned previously, a formulation of the vehicle
embodiment can be masticated with water to yield a compliant
material for placement into a tooth extraction socket during end
use either with or without an oxidized cellulose fibrous
reinforcement member, or it can be used in its non-hydrated form
preferably with a fibrous reinforcement member as illustrated in
Example 9. By contrast, the formulation described in Example 2
needs to be masticated with water to yield a compliant dough-like
material before it can be deployed during end use.
[0311] Regardless of which embodiment is deployed, it is important
that the formulation remain cohesively intact for as long as
possible following initial deployment, so as to enable the
formulation to 1) absorb fluids from the tooth extraction socket,
2) to gel with the fluids and to build its cohesive strength, and
3) to remain intact as a viable vehicle to facilitate controlled
release of active ingredients. A device which begins to erode and
disintegrate prior to gelation can lead to lower longevity during
use, so it can be appreciated that the best device is one that can
maintain its cohesive integrity for as long as possible under end
use conditions. In order to qualitatively assess these
characteristics, a static water soak test was devised for the
purpose of qualitatively testing each device's propensity to
swell/expand, or to disintegrate/dissolve under static conditions
vs. time. In this example, the three comparative samples from
Example 10, Formula 14C-2 with mineral oil, 12019-23-1 with
isopropyl palmitate, and 12019-23-2 with caprylic triglyceride,
were comparatively tested. The three comparative devices of similar
approximate weight are shown in FIG. 1. Each device was placed into
separate 11 ml glass vials with lids, and 2.5 g of water (pH
neutral distilled water, 20 degrees C.) was added to each as shown
in FIG. 2, with formulations from left to right including
12019-23-2, 12019-23-1, and 14C-2. The vial weights, the device
weights, and the added water weights were measured as follows:
Sample 12019-23-1 vial+lid=9.7394 g, tarred device wt.=0.6155 g,
and water weight=2.50 g; Sample 12019-23-2 vial+lid=10.2668 g,
tarred device wt.=0.6794 g, water weight=2.50 g; and Sample
12019-14C-2 vial+lid=9.8236 g, tarred device wt.=0.6745 g, and
water weight=2.50 g. The samples were then monitored vs. time
(FIGS. 3 through 6).
[0312] Visual inspection of the samples revealed that the relative
degree of swelling and disintegration was mirrored by the
qualitative compliance trends as recorded in Example 10. Namely,
the formula with higher compliance Formula 14C-2 tended to remain
cohesively intact and resisted delamination from its SO textile
members through the course of the experiment. By contrast, the
least compliant sample, 12019-23-2, exhibited the fastest relative
rate of swelling and disintegration, and it also exhibited evidence
of delamination from its SO textile member within the first 24
hours of the soak experiment.
[0313] These results do not necessarily imply that one type of oil
is better than another. Instead, it appears that the compliance and
cohesive strength characteristics of the device are important to
consider when formulating the device for longevity under static
soaking conditions. Based on the teachings of the prior examples,
the three oil types in the comparative samples could be formulated
in alternative ways to achieve optimum cohesive strength and
compliance characteristics. For example, one way to increase
compliance would be to increase the oil to wax ratio. Wax type can
also have an impact on compliance and cohesive strength. Another
way would be to decrease the volume fraction of particulates. Yet
another way would be to change the particle size distribution of
the PLGA and gelatin particulates. Another way would be to change
the knit density of the fibrous textile component in the
device.
[0314] Thus, when co-optimizing the device formulation and
construction for yielding the most desirable set of responses for
the end use application, including for example, tactile
characteristics during deployment, cohesive integrity, capacity to
absorb tooth extraction socket fluids, available concentrations of
active ingredients, active ingredient release rates; consideration
must be given not only to the chemical nature and ratios of the
vehicle components such as oil type, wax type, oil/wax ratio,
gelatin type, % of total dispersed solids within the oil/wax phase,
gelatin particle size distribution, PLGA particle size
distribution, ratio of the gelatin and PLGA dispersed ingredients,
but also to the macroscopic nature of the device's construction
such as knit density of the fibrous member, volume fraction of
fibrous members, surface chemistry of the fibrous members, degree
of oxidation of the fibrous members, and the resultant mechanical
properties of the fiber reinforced device).
Example 12. Relative Bupivacaine Release Characteristics from
Formulations Prepared with Hydrophilic and Hydrophobic
Components
[0315] The main purpose of this example is to illustrate the
variance in relative rates of bupivacaine (BUP) release that can be
achieved depending upon the hydrophobic or hydrophilic components
of the formulation, and on the morphological distribution of the
BUP active ingredient. Select versions of embodiments of the
formulation as described in Example 2 and the embodiments as
described in Examples 9 and 10 were prepared for comparative
purposes. Formulations were prepared in two ways: 1) using
bupivacaine free base encapsulated within PLGA microspheres; and 2)
using BUP free base that had been formulated directly into the
delivery device and not encapsulated by PLGA microspheres. Thus,
one of the primary differences among samples was the morphological
distribution of the BUP active ingredient inside versus outside of
the PLGA particles. A second primary difference was the relative
hydrophobicity of the formulation. The devices were immersed into
mildly acidic water (pH=2 prepared with HCl in deionized water) and
were incubated at 37 degrees C. for various lengths of time over a
24-hour interval for the hydrophilic devices, and over a 4-day
interval for the hydrophobic devices. Photos of the devices were
taken as a function of time to record their relative propensity to
either swell, disintegrate & dissolve, or to maintain cohesive
integrity under static soaking conditions as a function of time. In
addition, UV spectroscopy with specular beam detection was used to
follow the relative rate of bupivacaine release as a function of
time.
[0316] Importantly, the absorbance response from UV spectroscopy in
transmission mode with specular beam detection, as employed in this
example and in Example 13, is weighted by the presence of
molecular-scale components that have become solvated within the
liquid medium as opposed to components that have become dispersed
through erosion. This is relevant because it implies that soluble
components are preferentially detected, while dispersed components
are excluded from specular detection and can only be detected and
quantified via the use of an integrating sphere because they
scatter light diffusely. Thus, in order for one or more soluble
components to be detected in the supernatants of samples that have
been water-soaked under static conditions, molecular level
dissolution is a mandatory precursor. Moreover, in order for
dissolution to occur under static conditions, components would have
to first become inter-mixed with water at the molecular level
through a process that originates with molecular level diffusion.
Molecular level diffusion can occur via one or more of the
following pathways in any combination, including for example: 1)
water or other fluids entering the mother device; 2) active
ingredients or other components dissolving and egressing from the
mother device; 3) water diffusing into macro fragments that have
been eroded away from the mother device; 4) active ingredients or
other components leaching from macro fragments that have been
eroded away from the mother device; or 5) components egressing from
PLGA microspheres, including, microspheres that remain suspended
within the mother device, microspheres contained within macroscopic
fragments of the mother device, or microspheres that have become
freely dispersed within the supernatant water-phase.
[0317] Independent of the originating pathway, each of these
molecular-level processes requires translational motion of
molecular-scale entities across one or more concentration
gradients. By definition, concentration gradients will persist
under non-equilibrium conditions until the entire system comes to
equilibrium. In a closed system represented by a static water-soak
experiment, this implies that ingress and egress of molecules will
continue until the entire system reaches its equilibrium end point.
Macroscopic erosion is not a mandatory precursor for diffusion and
dissolution. However, if macro erosion does occur, it may indeed
lead to the faster appearance of molecular level entities that are
dissolved in solution, but dissolution is still the necessary
precursor for specular beam detection. Thus, when BUP is detected
in these experiments with UV spectroscopy, its detection is
evidence of its dissolution, which can only occur via diffusion and
dissolution from one of the pathways described above. Further, for
the case of BUP that originates from the interior of a PLGA
particle, it can only be detected if it has become dissolved,
necessitating that it must first migrate across one or more
concentration gradients represented by 1) the PLGA polymer that
constitutes the particle itself, where the interior of the particle
initially contains a higher BUP than the external chemical
environment; and 2) the matrix phase, which initially constitutes
the external chemical environment for a large fraction of the PLGA
particles that are dispersed therein.
[0318] A detailed accounting of sample compositions, experimental
details for sample preparations, measurement procedures, and
experimental results are provided below.
Hydrophilic Sample Compositions, Preparations, and Procedures.
[0319] Two comparative formulations comprising hydrophilic
components were based on compositions as discussed in Example 2.
Specifically, the two formulations for the present examples 918-1B
and 918-1i were prepared by using a formulation that was analogous
to that of formulas #4, #5, and #6 from DOE-DRAFT-6 containing 81%
microspheres by weight, 19% GLBG by weight, with GLBG gelatin as
the binder. This type of composition was previously observed to
provide a high relative BUP-dosage delivery that exhibited
acceptable compliance when masticated and hydrated with water.
However, for the purposes of the present example, the PLGA particle
size distribution was maintained at 100% 3-4 micron PLGA particles
as opposed to the distribution comprised of a mixture of small and
larger PLGA microspheres as described in DOE DRAFT-6 formula #7,
where a 10/90 w/w blend of distributions comprising D50=3.4 micron
and D50=42.1 micron particles was employed. Although from a
compliance perspective, a mixed PLGA particle size distribution
like that from formula #7 is one approach, a single PLGA particle
size distribution was employed in this example for facilitating
simple relative comparisons of cohesive integrity, release rates,
and relative compliance characteristics when comparing the two
formulations. The formulations comprising hydrophilic components
were prepared with Great Lakes Bovine Gelatin (GLBG), and with PLGA
microspheres that were made by SWRI using a solvent-borne Resomer
RG504 polymer with a spinning disc atomization drying process. The
PLGA microsphere samples had the following specifications: 1)
sample ID 18-0202-015-10 having an average particle size of D50=3.5
micron, and containing 20% by weight BUP free base; and 2) sample
ID 18-0202-015-7 placebo PLGA, also with an average particle size
of D50=3.5 micron, but with no BUP. The dry powder mixtures were
prepared using procedures outlined previously in Example 2. For the
case of sample 918-1i, BUP free base (Santa Cruz Biotechnology, CAS
#38396-39-3) was added directly to the dry powder mixture at a
level commensurate to the level used in sample 918-1B. The mixing
compositions of the comparative formulations are provided in Table
12-1 for the powders before and after hydration. The two
comparative formulations were designed to deliver a maximum dosage
of approximately 92 mg BUP per gram of hydrated device. Note that
the weight percentage of each ingredient was the same for each
device. The only difference was in the morphological distribution
of the BUP.
TABLE-US-00021 TABLE 12-1 Weight % compositions of hydrophilic
devices (dry powders before and after hydration). Calculations also
include the weight % concentration of BUP in the devices (before
and after hydration), and the effective available BUP concentration
for release during the pH-2 water-soak experiments. 918-1B dry
918-1B 918-1i dry 918-1i Ingredient powder hydrated gel powder
hydrated gel Great Lakes Bovine Gelatin 19.14% 10.93% 19.14% 10.93%
(GLBG) PLGA from 3.5-micron 64.69% 36.94% 0% 0% microspheres loaded
with 20% by wt. BUP free base Encapsulated BUP from 3.5- 16.17%
9.24% 0% 0% micron microspheres loaded with 20% by wt. BUP free
base PLGA from 3.5-micron 0% 0% 64.69% 36.94% placebo microspheres
BUP free base (non- 0% 0% 16.17% 9.24% encapsulated, directly added
to the vehicle) Water for hydration (pH-2, 0% 42.89% 0% 42.89%
dilute HCl) mg BUP/g device 162 92 162 92 Ratio of total water to
BUP N/A 58.3 N/A 58.3 (w/w) in water soak experiment Tarred Weight
(g) of N/A 1.1532 N/A 1.254 hydrated device added to 11 ml glass
vial Weight of water (g) in N/A 0.4946 N/A 0.5378 hydrated sample
Weight of additional water N/A 5.7186 N/A 6.2184 (g) added to 11 ml
vial Total water used in water N/A 6.2132 N/A 6.7562 soak
experiment (sum of water used for hydration + additional water that
was added to 11 ml glass vial) Effective Weight ratio of N/A 0.106
N/A 0.106 total device solids to water during the water-soak
experiment mg of available BUP per ml N/A 17.07 N/A 17.07 water
[0320] Before initiating the water-soak experiments, the dry
powders were first masticated with a fixed weight ratio of water
(pH-2, with dilute HCl added to deionized water) to gelatin of
3.92/1 w/w water/GLBG under ambient conditions (20 degrees C.) to
yield compliant gel-like mixtures. The samples were mixed with a
spatula in 15 ml HDPE beakers using procedures similar to those
reported in Example 2. The quantities of powders and water were
scaled to achieve a total masticated device weight that was
approximately 1.5 g in each case. During the mastication step, the
mixtures were initially observed to be creamy and low in viscosity.
After mastication, the beakers were covered with aluminum foil, and
the foil was removed at various times to check for gelation. Within
30 minutes, the samples had become solid and compliant gel-like
materials. At approximately 35 minutes after mixing, the gelled
devices were transferred and weighed into zero-tarred 11 ml glass
vials with lids. Next, at t=45 minutes after mixing, a specific
amount of pH-2 water was added to each vial, such that the total
water to BUP weight ratio was 7.0/0.12 w/w, inclusive of water that
was used during the mastication step. Thus, the samples were
allowed to gel for a total of 45 minutes after mixing prior to the
onset of the water soak experiment. Note that a constant water to
drug weight ratio of 7.0/0.12 was also used in each of the
comparative water-soak experiments for all of the samples, and the
same size vials were used to maintain similar surface to volume
ratios. For the present samples, this facilitated an equivalent net
reservoir of .about.17 mg BUP per ml water for potential elution
and delivery to the water phase throughout the course of the water
soak experiment. The vials were then incubated at 37 degrees C. for
the purposes of 1) tracking cohesive integrity vs. time (FIGS.
7a-7d, and 8) tracking the relative eluted drug concentration vs.
time via UV spectroscopy.
[0321] At t=1.5 hours, the vials were removed from the incubator
and a photo was taken. As illustrated in FIGS. 7a through 7d, both
samples had already started to swell and to disintegrate, but
sample 918-1i had already become noticeably more swollen and had
started to disintegrate to a higher degree than sample 918-1B. This
was particularly surprising in light of the fact that both samples
were formulated to have the same empirical composition (see Table
12-1), with the only difference being in the morphological
distribution of the BUP. The relative resistance of 918-1B to
erosion is believed to be a result of a synergy between the
plasticized polymer matrix phase and the PLGA-encapsulated BUP
microparticles that were dispersed therein, where microparticles of
this type appear to provide a type of mechanical reinforcement that
improves the cohesive integrity of the device.
[0322] In the next step, 2 ml of supernatant was removed from each
sample for UV analysis. The two glass vials were then closed and
were placed back into the incubator, and the two supernatant
aliquots were centrifuged at 3000 rpm for 5 minutes. Afterwards, 1
ml of each centrifuged liquid was used for UV absorption spectral
analyses. In this way, the relative level of dissolved BUP was
monitored as a function of time via UV absorbance intensities. When
the UV measurements were completed, the centrifuged aliquots of 2
ml in total for each sample were returned to their respective
vials, and the samples were allowed to continue incubating. This
sampling procedure was repeated at t=4 hours and again at t=24
hours after the onset of the water-soaking experiment.
UV Absorption Experiments of Samples 918-1i and 918-1B.
[0323] As mentioned above, aliquots were manually pipetted from the
top portion of the centrifuged supernatants, and 1 mL was loaded
into UV/VIS compatible cuvettes having outside dimensions=12
mm.times.12 mm, and inside path length=10 mm. For the case of
samples 918-1i and 918-1B, the net potential availability of BUP
for elution into the water phase was therefore approximately 17
mg/ml at maximum for the UV absorption experiments. Poly(methyl
methacrylate) cuvettes (Fisher brand) were used to limit cuvette
absorption within the range of detection for absorbance
measurements on a Tecan Infinite M200 Spectrometer within the range
of 2.30-1000 nm. Given the lack of absorbance at higher
wavelengths, spectrometer readings were typically measured between
250-350 nm. A wavelength step size of 2 nm with a bandwidth between
5-9 nm was used, and with 25 flashes, which was the number of
incident light exposure and detection occurrences that were signal
averaged at each wavelength. After each absorbance measurement, the
supernatants were collected and added back to the original glass
vials, such that the total volume in the elution experiment did not
change except for minor loss due to residual supernatant in the
pipette or UV cuvette.
[0324] Using the same instrument parameters and cuvettes, spectra
were also acquired for each individual ingredient in the mixtures
for the purpose of determining background contributions to the
overall absorbance spectra that were obtained for the fully
formulated mixtures. For the purposes of the background experiment,
the individual ingredients were either fully dissolved or were
fully dispersed in pH-2 water. In this way, the background
experiments were representative of the highest degree of spectral
background contribution that might be potentially observed for the
fully formulated devices if the devices were to completely
disintegrate and dissolve during the water-soak experiment.
[0325] The concentrations for these individual background
experiments were established from the effective ratio of each
individual ingredient to water that was used during the pH 2
water-soaking experiment on the fully formulated sample delivery
systems themselves. These concentrations, established from ratios
of values in Table 12-1, are reported in Table 12-2. The background
samples were aged for 24 hours under ambient conditions prior to
acquiring the UV spectra shown in FIG. 8. Inspection of these
background spectra revealed that the BUP itself was the strongest
chromophore in the mixture, and BUP was therefore the most
significant contributor to spectral absorption over the wavelength
range of interest (250-350 nm). Although the BUP free base has low
solubility in water, the mildly acidic conditions insured that the
BUP became protonated as the hydrochloride salt (BUP-HCl),
rendering it as completely soluble under these conditions. BUP-HCl
is known to be a strong chromophore with a documented UV absorption
maximum at 262 nm (Corciova, A., Eur. Chem. Bull., 2012, 2(8),
554-557).
[0326] As shown in FIG. 8, the spectral contributions from PLGA
placebo microspheres were negligible. However, a minor absorption
contribution was observed for bovine gelatin (GLBG), but this
contribution was observed to be minimal over the wavelength range
associated with BUP. Also, the PLGA-encapsulated BUP revealed
strong absorption after the 24-hour aging period, which was similar
in magnitude to that of the freely dissolved BUP itself. Thus, when
left unprotected by a matrix phase, the PLGA microspheres can
release enough BUP within 24 hours to completely saturate the
detector under the experimental conditions that were used in this
example. Although the absolute concentration of released BUP was
not measured in this experiment, it is important to note that the
relative amount that was released from the unprotected microspheres
was high enough to saturate the detector under the experimental
conditions associated with the water-soak experiment for the fully
formulated devices. This is noteworthy because when the BUP
encapsulated microspheres were protected by the matrix phase, the
net concentration of BUP-release after 24 hours was qualitatively
less than that exhibited by the unprotected microspheres. This
indicates that the matrix phases also play a substantial role in
mitigating diffusion.
[0327] The spectra acquired from the supernatants of samples 918-1i
and 918-1B are provided in FIG. 9. These results show that the
elution of BUP was faster when the BUP was morphologically
positioned to be outside of the PLGA microspheres as in sample
918-1i. By contrast, the elution of BUP was deterred by PLGA
encapsulation in sample 918-1B. In other words, the system
containing BUP that was encapsulated within PLGA microspheres
(918-1B) was observed to release BUP more slowly than the system
that contained BUP that was directly formulated into the vehicle
(91.8-1i). Thus, when PLGA was used to encapsulate the active
ingredient, its rate of release into solution was attenuated. Note
that a similar trend was observed for analogous delivery made with
hydrophobic components, but the release rates were further
attenuated by the hydrophobicity of the vehicles (described
below).
[0328] At t=4 hours, the level of BUP that was released from 918-1i
was already at a high enough level to saturate the detector. By
contrast, the level of BUP released from 918-1B was lower, and it
was still within the range of instrumental detection. However, the
amount of BUP that was released from both samples was high enough
within 24 hours to saturate the UV detector.
[0329] In a separate experiment, a wavelength of 270 nm was chosen
for establishing a separate calibration curve for the BUP
concentration versus absorbance in pH-2 water. That is, the weight
of BUP free base in mg/ml in dilute HCl was plotted versus
absorbance at wavelength=270 nm. This calibration equation provided
in Table 12-3 was used to roughly estimate the concentration of
dissolved BUP as a function of time during the water-soak
experiment for the fully formulated samples. Note that since these
were single beam acquisitions with no reference cuvette, a separate
single beam absorption spectrum from pH-2 water was subtracted from
the BUP sample spectra before establishing the correlation.
[0330] Before estimating the effective elution concentration of BUP
from the fully formulated sample mixtures, the absorbance values at
270 nm were first corrected by subtracting an absorbance
contribution from fully dissolved GLBG. This actually represented
an overcorrection since the GLBG did not become fully dissolved
during the water-soak experiments on the hydrated samples. Thus,
whenever this overcorrection resulted in a negative value during
early periods of the water-soak experiment, the BUP estimate was
equated to zero. For the case of samples 918-1i and 918-1B, the UV
absorbance intensity of GLBG that was dissolved in water at a
concentration of 0.0203 g per g pH-2 water was used to make this
absorbance correction at a concentration equivalent to the net
concentration of gelatin that was present and available for
complete dissolution during the water-soak experiment on the
devices. Note that no correction was made for the presence of PLGA
since the separate UV experiments revealed that absorbance
contribution of PLGA was negligible within the wavelength range
attributable to BUP absorption.
[0331] Note also that the detector becomes signal-saturated at
absorbance values approaching 4 absorbance units. Through the
course of the water-soaking experiments, this saturated detector
condition was eventually achieved for each sample. For the purposes
of the present example, the [BUP] calibration curve was used to
estimate BUP elution concentrations only for cases where the
absorbance was <3.9. When the detector saturation level was
reached, the estimated BUP elution concentration was reported as
equal to or greater than the value calculated from the calibration
plot. Note that successive dilutions could be used to bring the
absorbance values back within the detection range, but for the
purposes of this example, these additional experiments were not
necessary to illustrate the important differences among the sample
types. In the next step, the estimated BUP elution concentration
[BUP].sub.t was ratioed against the total theoretical concentration
[BUP].sub.theoretical to allow for qualitative comparison of
relative elution rates among the devices from the present
example.
TABLE-US-00022 TABLE 12-2 Table entries include concentrations of
individual ingredients in pH-2 water (w/w) for acquisitions of
individual background UV spectra shown in 8; weights of device
mixtures per g of pH-2 water during the water-soak experiment;
absorbance contributions of individual ingredients at 270 nm;
absorption at 270 nm of supernatants from devices during the water-
soak experiment; GLBG overcorrected absorption values; estimated
[BUP] at each time interval (from calibration equation in Table
12-3); and the estimated fraction of eluted BUP based on an initial
theoretical concentration of BUP that was available from the device
(i.e., 17.07 mg/ml for the hydrophilic devices as reported in Table
12-1). Note that when an absorbance overcorrection resulted in a
negative value, the value was denoted as zero (marked with an
asterisk). When the measured absorbance values were approaching the
detector limit, the effective BUP concentration was denoted as
equal to or less than the theoretical maximum of ~17 mg/ml (also
denoted with an asterisk). Relative Relative Relative Absorbance at
Absorbance at Absorbance at Weight (g) per 270 nm at t = 270 nm at
t = 270 nm at t = Ingredient gram pH-2 water 1.5 hrs. 4 hrs. 24
hrs. pH 2 dilute HCl 1 N/A N/A 0.2818 PLGA Placebo 0.0686 N/A N/A
0.3069 Microspheres in pH 2 (dispersed) water GLBG fully dissolved
0.0203 N/A N/A 1.3871 in pH 2 water (fully dissolved) BUP free base
(fully 0.0171 N/A N/A > or = 4 (i.e., dissolved) saturated
signal) 918-1B 0.106 0.6703 2.504 3.8087 918-1B (corrected for --
~0* 1.1169 2.4216 GLBG) 918-1B estimated -- ~0* 1.4 *Between 3 and
17.1 BUP elution concentration [BUP]t using equation from Table
12-3 (mg/ml) 918-1B estimated -- ~0* 0.08 > or = 0.17* fraction
of BUP elution = [BUP]t/ [BUP]theoretical; [BUP]theor. = 17.1
918-1i 0.106 2.1606 3.7002 3.6891 918-1i (corrected for -- 0.7735
2.3131 2.3020 GLBG) 918-1i estimated BUP -- 0.96 2.9 *Between 3 and
17.1 elution concentration [BUP]t using equation from Table 12-3
(mg/ml) 918-1i estimated -- 0.06 0.17 > or = 0.17* fraction of
BUP elution = [BUP]t/ [BUP]theoretical; [BUP]theor. = 17.1
TABLE-US-00023 TABLE 12-3 BUP calibration equation as obtained from
a linear best fit of absorption at 270 nm vs. BUP concentration
(mg/ml) in pH-2 water. This table provides the absorbance intensity
for BUP free base that was fully dissolved in pH-2 water over the
detectable range of [BUP], expressed in mg/ml. Note that the
absorbance values as reported below represent corrected values that
were obtained by subtracting a single-beam absorbance spectrum of
pH-2 water from the single-beam absorbance spectra of the BUP
samples. [BUP] mg/ml Relative Absorbance Intensity 2.2364 2.7859
1.2300 1.6119 0.72684 0.9195 0.22364 0.2290 0.022364 -0.0248
0.0022364 -0.0511 0.00022364 -0.0394 0 0 R.sup.2= 0.998 Slope=
1.2787 y-intercept= -0.0310
Hydrophobic Sample Compositions, Preparations, and Procedures.
[0332] The comparative formulations comprising hydrophobic
components were based on compositions as discussed in Examples 9
and 10. Specifically, three formulations were prepared for this
example, 14C-3A, 14C-3B2, and 14C-3A Placebo, by using a
formulation that was similar to that of Formula 14C-2 with a few
exceptions, including: 1) the 30/70 w/w blend of 5 micron to 41
micron PLGA placebo microspheres were substituted with
distributions comprised of 100% of smaller sized PLGA microspheres;
2) 14C-3A was formulated with PLGA microspheres containing 20% BUP
free base by weight (SWRI, 20% BUP free base loaded PLGA
microspheres based on Resomer RG504, prepared using a using a
spinning disc spray-dry atomization process, sample ID
18-0202-015-p21; D50=4.3 microns; photo provided in FIG. 7e); 3)
14C-3B2 was formulated with BUP free base (Santa Cruz
Biotechnology, CAS #38396-39-3) that was added directly to the
formulation together with PLGA placebo microspheres (SWRI, PLGA
placebo microspheres based on Resomer RG504, prepared using a using
a spinning disc spray-dry atomization process, sample ID
18-0202-105-15; D50=5 microns); and 4) the compositions of 14C-3B2
and 14C-3A Placebo were subtly modified to achieve compliance
characteristics suitable for textile impregnation.
[0333] The formulations prior to textile impregnation were prepared
using the same materials as those used in Example 9. A gelatinous
melt-recrystallized amalgam of mineral oil (MO) and beeswax (BW)
was prepared at the same ratio as was used previously for 14C-2
(5/1 w/w oil to wax). In the next step, Great Lakes bovine gelatin
(GLBG) was added to the amalgam to form a premix. For the cases of
14C-3A and 14C-3B2, the same weight ratios were used for the premix
as reported for 14C-2 from Example 9 (55.10 weight % MO, 11.10
weight % BW, and 33.39 weight % GLBG). For the case of 14C-3A
Placebo, a slightly lower level of GLBG was used in the premix for
the purpose of adjusting the viscosity.
[0334] In the next step, PLGA microspheres were dispersed into the
premixtures of MO/BW/GLBG using the same procedures as reported in
Example 9. The 14C-3A formulation was made by dispersing the 20%
BUP-loaded PLGA microspheres (D50=4.3 microns) into the premixed
vehicle under ambient conditions. The resulting dispersion had very
similar compliance and viscosity characteristics to sample 14C-2 as
was made previously, in spite of the use of the smaller PLGA
particle size distribution. The 14C-3B2 formula was similarly
prepared, but in the first step, BUP free base was dispersed
directly into the premix of MO/BW/GLBG under ambient conditions,
and then the PLGA placebo microspheres were added (D50=5 microns).
In a first attempt, the 14C-3B2 formula was targeted to have an
identical composition to that of 14C-2. However, the total
percentage of dispersed solids were initially too high to yield a
compliant dispersion and a dry blend was formed instead. For this
reason, the total weight % of dispersed solids, predominantly BUP
free base and PLGA placebo microspheres in the case of 14C-3B2,
were reduced to a level that allowed for the formulation of a
dispersion that would be compliant enough to impregnate a fibrous
textile. Finally, a third formula, 14C-3A Placebo, was also
prepared, analogous to 14C-3A and 14C-3B2, but containing PLGA
placebo microspheres (SWRI, PLGA placebo microspheres based on
Resomer RG504, prepared using a using a spinning disc spray-dry
atomization process, sample ID 18-0202-105-15; D50=5 microns)
instead of BUP loaded microspheres. The 14C-3A Placebo formula,
like the 14C-3B2 formula, also required a slightly lower weight
percent of dispersed solids to achieve compliance characteristics
that were suitable for textile impregnation. In this case, the
adjustment was made by diluting the vehicle with the 5/1 (w/w)
amalgam of melt-recrystallized MO and BW. The three resulting
sample formulations exhibited qualitatively similar compliance
characteristics to one another, and each was used to prepare
cellulose textile-impregnated devices for the purposes of the
present example using procedures as reported in Examples 9 and 10.
The mixing compositions of the comparative formulations are
provided in Table 12-4.
TABLE-US-00024 TABLE 12-4 Weight % compositions of hydrophobic
vehicles for use in preparing textile-impregnated devices.
Calculations also include the net weight % concentration of BUP in
each vehicle, the net PLGA polymer weight % (i.e., ~80% of the
weight of BUP loaded microspheres, and 100% of placebo
microspheres), and the total weight % of dispersed solids. 14C-3A
Vehicle Mixture Composition 14C-3A 14C-3B2 Placebo Mineral Oil
23.03% 28.39% 28.12% Beeswax 4.61% 5.68% 5.62% Bovine Gelatin
13.85% 17.07% 12.68% 5 um PLGA Placebo 0% 39.09% 53.57%
microspheres 4.3 micron 20% BUP free base 58.51% 0% 0% loaded PLGA
microspheres BUP free base (directly added 0% 9.77% 0% to vehicle)
TOTAL 100.00% .sup. 100% .sup. 100% Total BUP in vehicle 11.70%
9.77% 0% Total PLGA polymer in Vehicle 46.81% 39.09% 53.57% Total %
dispersed solids in 76.97% 71.61% 71.88% vehicle
Impregnation of Cellulose Fiber Textiles.
[0335] The delivery devices were prepared using procedures as
reported in Example 9. The construction used for the devices was
like that reported for set-2 in Table 5 of Example 9, with two
orthogonal Surgicel Original (SO) textiles having dimensions of
approximately 1.8.times.3.8 cm each. The average single SO textile
weight was 0.0470 g (n=35, SD=+/-0.0016 g), resulting in an average
cellulose textile weight contribution of 0.094 g per device. The
weight of each impregnated vehicle ranged from approximately 0.73 g
to 0.75 g per device. The device compositions are reported in Table
12-5. The comparative sample formulations were designed to deliver
a maximum BUP concentration of approximately 104 mg BUP per g of
the 14C-3A device and 88 mg per g of the 14C-3B2 device.
TABLE-US-00025 TABLE 12-5 Weight % compositions of hydrophobic
textile-impregnated devices. The vehicle compositions as reported
in Table 12-4 were impregnated into two orthogonally oriented SO
textiles. The calculations for compositions also include the
concentration of BUP, and the effective available BUP concentration
for release during the water-soak experiments. Note that when the
devices were transferred to 11 ml glass vials, a small amount of
vehicle weight was lost. This loss was taken into account to insure
that the correct water to BUP weight ratios were employed during
the water-soak experiment (i.e., to achieve a BUP reservoir
concentration of approximately 17 mg BUP/ml water, which was the
same concentration that was used during the water-soak experiments
on the comparative devices that were made with hydrophilic
components). Ingredient 14C-3A 14C-3B2 14C-3A Placebo Great Lakes
Bovine Gelatin 12.35% 15.36% 11.29% (GLBG) Mineral Oil (MO) 20.53%
25.54% 25.04% Beeswax (BW) 4.11% 5.11% 5.01% PLGA polymer (i.e.,
representing 41.73% 0% 0% 80% of the weight of 4.3-micron
microspheres loaded with 20% by wt. BUP) Encapsulated BUP (i.e.,
10.43% 0% 0% representing 20% by weight of the 4.3-micron
microspheres loaded with 20% by wt. BUP) PLGA polymer from 5-micron
0% 35.16% 47.7% placebo microspheres BUP free base
(non-encapsulated, 0% 8.79% 0% directly added to the vehicle) SO
textiles 10.84% 10.05% 10.96% mg BUP/g device 104 88 0 Target ratio
of pH 2 water to BUP 58.33 58.33 N/A (w/w) in water soak experiment
Weight of Device as made (g) 0.8296 0.8257 0.8457 Weight of Vehicle
as made (g) 0.7372 0.7304 0.7529 Tarred Weight (g) of device added
0.8176 0.8250 0.8290 to 11 ml glass vial Weight of vehicle after
transfer to 0.7252 0.7297 0.7363 vial (g) (containing (containing
11.70% 9.77% BUP) BUP) Weight of pH 2 water (g) added 4.9522 4.1594
5.0071 to 11 ml vial mg of available BUP per ml pH-2 17.13 17.13 0
water
[0336] When the devices comprising the textile-impregnated
formulations were transferred to the 11 ml glass vials for the
water soak experiment, the level of pH-2 water was adjusted to
achieve a net level of approximately 17 mg BUP per ml of water as
the theoretical maximum level of available BUP if all of the BUP
were to be released and dissolved into the water phase. Note that a
constant water to drug weight ratio of 7.0/0.12 was also used in
each of the comparative water-soak experiments for the previously
described comparative samples that were prepared with hydrophilic
components. Thus, independent of the device type, experimental
conditions were established to allow for an equivalent net
reservoir of 17 mg BUP per ml water for potential elution and
delivery to the water phase throughout the course of the water soak
experiments. The vials were incubated at 37 degrees C. for the
purpose of tracking cohesive integrity vs. time (FIGS. 10a through
10e), and for tracking the relative BUP release concentration
versus time via UV spectroscopy (i.e., the relative level of BUP
that was released and dissolved in the supernatants).
[0337] As illustrated in FIGS. 10b, 10c, and 10d, there were no
major visual differences among the supernatants of samples at t=1.5
hours, t=4 hours, and at t=24 hours. Unlike the formulations
comprising hydrophilic components, there was no evidence of haze in
the supernatants from disintegration of the vehicles. This behavior
was observed to continue throughout the 4-day course of the
water-soak experiment. The behavior was surprising from the
standpoint that even in the absence of visual disintegration,
measurable relative concentrations of soluble BUP were still
released. At t=24 hours, and especially after 4 days, there was a
subtle degree of swelling of sample 14C-3B2, but the degree was
relatively minimal when compared to the degree of swelling and
disintegration that had occurred in both samples 918-1i and 918-1B
after only one day.
[0338] When the vials were removed from the incubator, 1 ml
aliquots of supernatant were removed from each sample for UV
analysis. The glass vials were then closed and were placed back
into the incubator to continue incubating through the course of
time required to complete the UV absorption measurements. Unlike
the supernatants from the earlier samples, the supernatant aliquots
from the samples comprising hydrophobic components were not
centrifuged. This was because no visual erosion and disintegration
had occurred among the samples as evidenced by the lack of haze
owing to the lack of water-dispersed solids. Again, each of the 1
ml aliquots was used for UV absorption spectral analyses, and the
relative levels of dissolved BUP were monitored as a function of
time for each of the supernatant samples. When each of the UV
measurements was completed, the 1 ml aliquots for each sample were
returned to their respective vials, and the samples were returned
to continue incubating at 37 degrees C. This sampling procedure was
repeated at t=4 hours, t=24 hours, and at t=96 hours after the
onset of the water-soaking experiment.
[0339] Importantly, the longer duration of the water-soak
experiment for samples 14C-3A, 14C-3B2, and 14C-3A Placebo in
comparison to samples 918-1i and 918-1B was made possible not only
because of slower BUP release rates, but also because of minimal
swelling and minimal disintegration among the samples. The samples
that were made with hydrophobic components were observed to
maintain their mechanical cohesive integrity for significantly
longer periods of time while soaking under static conditions than
their comparative counterparts that were made with hydrophilic
components. The unanticipated benefit of this behavior includes the
potential to create formulations for longer term use in the end use
application than would otherwise be possible with the comparable
formulations.
[0340] It is also important to note that the sample formulations
comprising hydrophobic components contained a higher weight % of
BUP per unit device weight than their hydrated counterpart
formulations comprising hydrophilic components. As noted in prior
examples, the latter require hydration prior to deployment to form
compliant dough-like substances to render them as suitable for
clinical use in the end use application. The addition of water
during the hydration step results in an unavoidable dilution of the
available BUP dosage per unit weight. By contrast, the comparative
samples devices do not require hydration because they are
formulated to have the necessary compliance needed for clinical use
in the end use application. Consequently, the net delivery dosage
of BUP per unit weight can be adjusted to higher levels in the
formulations with hydrophobic components when compared to the
formulations with hydrophilic counterparts. Moreover, the
differential in maximum dosage is similar when volume is taken into
consideration. In a volume-restricted application, as is the case
for an oral tooth socket cavity, the higher active ingredient
dosage per unit weight of a formulation with hydrophobic
ingredients translates to a higher delivery dosage of BUP per unit
volume than would otherwise be possible with a comparable device.
This unanticipated benefit provides an expanded opportunity to
create formulations with higher net dosage delivery levels for use
over protracted periods of time during end use if so desired.
UV Absorption Experiments of Hydrophobic Samples.
[0341] Instrument parameters and procedures were the same as those
used above, but with one major exception. Namely, as discussed
earlier, the 1 ml aliquots were pipetted directly from the
supernatants and then were analyzed without centrifuging. The 1 mL
aliquots were then loaded into the same types of UV/VIS compatible
cuvettes (outside dimensions=12 mm.times.12 mm, inside path
length=10 mm). Again, the net potential availability of BUP for
elution into the water phase was approximately 17 mg/ml at maximum
for the duration of UV absorption experiments. Spectrometer
readings were measured between 250-350 nm. A wavelength step size
of 2 nm with a bandwidth between 5-9 nm was used, and with 25
flashes, the number of incident light exposure & detection
occurrences that were signal averaged at each wavelength. After
each absorbance measurement, the supernatant was collected and was
then added back to the original glass vial, such that total volume
in the elution experiment did not change except for minor loss due
to residual supernatant in the pipette or UV cuvette.
[0342] FIGS. 11a through 11d provide four relative absorbance vs.
wavelength plots for the supernatants of samples 14C-3A, 14C-3B2
and 14C-3A Placebo with each plot delineating a separate soaking
time point, including points at t=1.5 hours, t=4 hours, t=24 hours,
and at t=96 hours after the onset of the water soaking experiments
in pH 2 water. As was the case for the comparative samples, these
plots reveal that BUP was released more slowly when the BUP was
encapsulated within PLGA in 14C-3A, and more quickly when the BUP
was formulated directly into the formulation as in sample 14C-3B2.
All three delivery systems showed evidence of component dissolution
as a function of time with the wavelengths between approximately
260-270 nm providing the clearest visual delineation of the
absorbance differences among the supernatants at the various time
points. Although the partial dissolution of other ingredients may
contribute to absorption in this range (as evidenced by the
placebo), BUP-HCl is known to be a strong chromophore with a
reported UV absorbance maximum of 262 nm (Corciova, A., Eur. Chem.
Bull., 2012, 2(8), 554-557). Hence, the evolution of absorption in
the 260-270 nm range is strongly influenced by the protonation of
BUP and by its subsequent dissolution vs. time. Thus, FIGS. 11a
through 11d collectively reveal that BUP was released more slowly
from the delivery system when the BUP was encapsulated within PLGA
(formulation 14C-3A), and more quickly when the BUP was formulated
directly into the hydrophobic vehicle (formulation 14C-3B2).
[0343] FIGS. 12a, 12b, and 12c provide an alternative
representation of the same data. Specifically, three plots are
provided with each plot representing the time evolution of
absorption curves for each individual sample. The plot representing
the time evolution of the placebo device, 14C-3A Placebo without
BUP, illustrates the egress of water-soluble components from the
formulation itself as a function of time, components other than
BUP, such as GLBG and SO. These types of components would be
expected to contribute to the overall background absorption from
supernatants of fully formulated samples that contain BUP.
Regarding the devices with BUP, the growth in absorption intensity
vs. time can be seen over the wavelength range 260-270 nm, with
faster growth occurring from the device where BUP was formulated
directly into sample 14C-3B2. The increase in absorption intensity
from conversion of BUP free base to the soluble amine hydrochloride
(BUP-HCl) was paralleled by a measurable increase in the pH of the
supernatants for the samples containing BUP. The following pH
values were measured after 10 days of soaking in pH-2 water at 37
degrees C.: 14C-3A (BUP free base inside microspheres)=3.51;
14C-3B2 (BUP free base formulated directly into the vehicle and
outside of the microspheres)=3.90; and 14C-3A Placebo (no BUP): pH
of solution=2.18. The minimal change in pH for the placebo
indicates that the other soluble components that contribute to UV
absorption in the 260-270 nm range have minimal impact on pH when
compared to the effect of the BUP free base itself. Moreover, the
highest degree of acid neutralization occurred when the BUP free
base was formulated directly into the sample, a result that
corroborates with faster release as measured by UV
spectroscopy.
[0344] Using the reported UV absorbance maximum for BUP-HCl (262
nm), FIG. 13 provides the evolution of the absorbance intensity as
a function of time for each of the device types. This plot is
presented with a natural log time scale to better illustrate the
large differences in the rates of egress among the device types.
For illustration purposes, each of the data sets were empirically
fit to a simple exponential growth function. The function and best
fit parameters are provided in Table 1.2-6. These trends illustrate
the relative difference in release rates afforded by 1) the
morphological distribution of the BUP, and by 2) the relative
hydrophobicity of the formulation. For example, the trends reveal
that the relative release rate of BUP increases with the use of
hydrophilic components, and then decreases when the BUP is
encapsulated by PLGA. Specifically, the relative BUP release rate
was observed to trend as follows:
918-1i>918-1B.about.14C-3B2>14C-3A.
[0345] These trends also reveal that independent of the other
ingredients, PLGA encapsulation attenuates the relative BUP release
rate. Surprisingly however, the initial relative release rate from
14C-3B2 containing BUP that is formulated within the sample without
PLGA encapsulation was observed to be approximately the same as
that of 918-1B in which the BUP was encapsulated within
microspheres. This counterintuitive result reinforces that the
present formulation affords the opportunity to control release
rates by virtue of employing multiple factors, either alone or in
any combination, including the use of BUP encapsulated by PLGA, the
use of freely formulated BUP, and the use of formulations with
varying degrees of hydrophobic and hydrophilic ingredients. For
example, if a short time-duration release profile is desired, 1-3
days, then a formulation with hydrophobic ingredients containing
BUP without PLGA encapsulation can be used to achieve similar
results to those of a formulation with hydrophilic ingredients
containing BUP that is encapsulated within PLGA microspheres.
TABLE-US-00026 TABLE 12-6 Exponential growth function along with
the adjustable parameters that were used to achieve an iterative
best fit of the relative absorbance intensity at 262 nm vs. time
data for the supernatants of the hydrophilic and hydrophobic
devices while soaking in pH 2 water at 37 degrees C. A plot of the
best fit data is provided in FIG. 13. Data were fit to the
following functional form: Abs = C1 + C2*[1 - exp(-C3*time)], where
Abs = absorbance at 262 nm; C1, C2, and C3 are constants derived
from the iterative best fit of the data; and time = soak time in
hours. The adjustable parameters are provided below. Sample C1 C2
C3 R.sup.2 918-1i (fastest rate) -0.53 4.29 1.01 .99 918-18 -2.68
6.44 0.57 .99 14C-3B2 -2.22 5.90 0.55 .99 14C-3A 0.56 10.36 0.003
.99 14C-3A Placebo 0.80 1.36 0.019 .97
[0346] The [BUP] calibration line from Table 12-3 (i.e., of [BUP]
mg/ml vs. relative absorbance intensity at 270 nm) was used to
estimate the approximate BUP concentration in the supernatants as a
function of time during the pH 2 water soak experiment. However,
before estimating the effective elution concentration of BUP from
samples 14C-3A and 14C-3B2, the supernatant absorbance values at
270 nm were first corrected by subtracting the absorbance
contribution from the 14C-3A Placebo sample, which in essence
equates to a mixed contribution of possible absorbances from GLBG,
SO, and perhaps even from PLGA, MO, and BW in the absence of BUP.
In this way, the corrected absorption values provided an estimate
of the relative BUP absorbance contribution, in this case soluble
BUP-HCl. These data and calculations are provided in Table
12-7.
[0347] The estimates of [BUP] versus time were limited to
supernatants with net absorbance values of less than about 3.9,
below the detector saturation level. When the detector saturation
level was reached, the estimated BUP elution concentration was
reported as equal to or greater than the value calculated from the
calibration line in Table 12-3, or in other words, greater than or
equal to the calculated BUP concentration at an absorbance value
approaching 3.9 but less than the maximum value of
[BUP].sub.theoretical. Note that successive dilutions could have
been used to bring the absorbance values back within the detection
range, but for the purposes of this example, these additional
experiments were not necessary in order to illustrate the important
differences among the sample types. In the next step, the estimated
BUP elution concentration [BUP]t was ratioed against the total
theoretical concentration [BUP].sub.theoretical thereby allowing
for comparison of relative BUP elution rates among the various
samples.
TABLE-US-00027 TABLE 12-7 Relative absorption of supernatants from
hydrophobic devices at 270 nm vs. time (hrs.) during the pH 2
water-soak experiment, including the estimated [BUP] at each time
interval, and the estimated fraction of eluted BUP based on an
initial theoretical concentration of BUP that was available from
the device (i.e., 17.13 mg/ml for the hydrophobic devices as
reported in Table 12-5). The absorption from the devices containing
BUP were corrected for background contributions from non-BUP
ingredients that may have dissolved (e.g., SO, GLBG) or dispersed
(e.g., PLGA, MO, BW) as a function of time during the soak
experiment. These contributions were roughly estimated from the
absorbance of the 14C-3A Placebo at 270 nm. Note that when the
absorbance correction resulted in a negative value, the correction
was denoted as zero (marked with an asterisk). When measured
absorbance values were at or approaching the saturation point of
the detector, the effective BUP concentration was denoted as
greater than the calculated value, but less than the theoretical
maximum of ~17 mg/ml (also denoted with an asterisk). Weight (g)
270 nm 270 nm 270 nm 270 nm per ml pH 2 Abs t = Abs t = Abs t = Abs
t = Device water 1.5 hrs. 4 hrs. 24 hrs. 96 hrs. 14C-3A Placebo
0.8290 0.4997 0.7690 1.0712 1.7985 14C-3A 0.8176 0.3296 0.4904
1.0345 2.8987 14C-3A (corrected for -- ~0* ~0* ~0* 1.1022
background contributions) 14C-3A estimated BUP elution -- ~0* ~0*
~0* 1.4 concentration [BUP]t using equation from Table 12-3 (mg/ml)
14C-3A estimated fraction of -- ~0* ~0* ~0* 0.08 BUP elution =
[BUP]t/ [BUP]theoretical; [BUP]theor. = 17.1 14C-3B2 0.8250 0.7817
2.6323 3.4235 3.6308 14C-3B2 (corrected for -- 0.2820 1.8633 2.3523
1.8323 background contributions) 14C-3B2 estimated BUP elution --
0.33 2.35 2.98 Between concentration [BUP]t using 2.3 and 17.1
equation from Table 12-3 (mg/ml) 14C-3B2 estimated fraction of --
0.02 0.14 0.17 = or >0.14 BUP elution = [BUP]t/
[BUP]theoretical; [BUP]theor. = 17.1
Summary of Results.
[0348] The cohesive integrity under static soaking conditions, as
qualitatively gauged by the relative degrees of visual swelling,
device disintegration, and haze in the supernatants from
disintegrated material, was qualitatively observed to improve with
the use of hydrophobic ingredients. The following trend was
observed from lowest to highest relative degree of visual swelling,
disintegration, and development of supernatant haze: 14C-3A
(hydrophobic vehicle, BUP encapsulated within PLGA
microspheres)<14C-3B2 (hydrophobic vehicle, BUP formulated
directly into vehicle containing placebo PLGA
microspheres)<<918-1B (hydrophilic vehicle, BUP encapsulated
within PLGA microspheres)<918-1i (hydrophilic vehicle, BUP
formulated directly into vehicle containing placebo PLGA
microspheres). The relative release rate of BUP was observed to
increase with the use of hydrophilic ingredients, and with BUP that
was not encapsulated: 918-1i>918-1B.about.14C-3B2>14C-3A.
[0349] The trends reveal that from the standpoints of mechanical
cohesive integrity and relative release rates, the hydrophobic
ingredients are best suited for formulations wherein the intention
is to achieve longer-term usage in the end application. For
example, the more hydrophilic formulations disintegrate more
quickly under static conditions, rendering them most useful for
shorter-term end use durations. By contrast, the more hydrophobic
formulations, particularly with fiber reinforcement, provide
longer-term cohesive integrity under static soaking conditions,
which render them as well-suited for both short-term and
longer-term use. The relative BUP release rates also corroborate
with these conclusions. Namely, the hydrophilic formulations afford
faster release than the hydrophobic formulations. Thus, one lever
that is useful in preparing a formulation with a controlled release
profile is the relative hydrophobicity of the formulation, where
the more hydrophobic the formulation, the slower the release. A
second lever that has proven to be useful for preparing devices
with controlled BUP release is PLGA encapsulation of BUP.
[0350] Independent of other features of the formulation, PLGA
encapsulation was observed to attenuate the relative BUP release
rate. Surprisingly however, the relative release rate from a more
hydrophobic formulation containing BUP without PLGA encapsulation
was observed to be similar to that of a more hydrophilic
formulation wherein the BUP was encapsulated within PLGA
microspheres. This result reinforces that the present formulation
technology affords the opportunity to control release rates by
virtue of employing multiple factors, either alone or in
combination, including 1) using PLGA to encapsulate BUP; 2)
adjusting the degree of hydrophobicity; and 3) incorporating BUP
with no PLGA encapsulation. For example, if a short time-duration
release profile is desired of 1 to 2 days, then a more hydrophobic
formulation containing BUP without PLGA encapsulation can be used
to achieve similar results to a more hydrophilic formulation
containing BUP that is encapsulated within PLGA microspheres. In
another example, microencapsulation of active ingredients can be
used in combination with free, non-encapsulated active ingredients,
to create devices exhibiting an adjustable range of relative BUP
release rates, depending on the ratio of encapsulated BUP to free
BUP. Moreover, when a more hydrophobic formulation is employed,
even higher net delivery dosages per unit weight device can be
achieved if so desired as further demonstrated in Example 13.
Example 13. Formulations Compromising Hydrophobic Ingredients and
Containing Mixtures of Encapsulated and Non-Encapsulated
Ingredients
[0351] More hydrophobic formulations like those described in
Example 12 were prepared for this example. The objectives were to
demonstrate various methods that can be used to control BUP release
from a more hydrophobic formulation, to demonstrate methods by
which the maximum dosage level of BUP can be raised to even higher
levels, and to demonstrate formulation flexibility that allows for
the incorporation of additional dispersed ingredients, such as pH
modulators, without adversely affecting rheological characteristics
and release characteristics.
[0352] Factors for this experiment included: 1) use of
non-encapsulated BUP-free base; 2) use of PLGA-encapsulated
BUP-free base; 3) the use of mixtures of PLGA-encapsulated BUP-free
base and non-encapsulated BUP-free base; and 4) the use of pH
modulators citric acid (Sigma-Aldrich, cat. #251275, CAS #77-92-9)
and dibasic sodium citrate sesquihydrate (Sigma-Aldrich cat.
#71635, CAS #6132-05-4, referred to herein as sodium citrate).
These experiments demonstrate the use of combinations of
encapsulated and non-encapsulated BUP to create devices with even
higher possible dosage loadings of BUP or other active ingredients
if so desired. Moreover, the experiments demonstrate that a range
of release rates are possible depending upon the ratio of the
encapsulated to non-encapsulated ingredients.
[0353] Using procedures outlined in Example 12, compositions were
mixed and were used to impregnate two orthogonally arranged
Surgicel Original (SO) textiles for the purpose of forming
control-release delivery devices. The compositions of the
formulations and devices are provided in Tables 13-1 and 13-2,
respectively. The devices were then subjected to pH-2 water-soak
testing at 37 degrees C. and, using methodology similar to that
which was described in Example 12, UV spectroscopy was used to
estimate the relative concentration of BUP that had diffused or
eluted into the supernatants as a function of time.
[0354] The concentration of BUP at each time increment was
estimated by using a two-step procedure. First, the UV absorption
values from the supernatants of the pH-2 water-soaked devices were
background corrected by subtracting the UV spectra of the
supernatant from a water-soaked placebo device which was used to
approximate the absorption contributions from non-BUP components,
sample 14C-3E Placebo in this example. FIG. 14 displays a relative
absorbance vs. time comparison of placebo devices 14C-3E (with
citric acid) and 14C-3A (without citric acid). Absorbance vs. time
data show that although the devices produced soluble components in
the absorbance region overlapping with BUP, there was no
significant effect of citric acid under pH-2 soaking conditions.
Consequently, to simplify analyses, the 14C-3E absorbance data were
used for all pertinent spectral background corrections relating to
the spectral absorbance of devices containing BUP.
[0355] In the second step, the background-corrected absorption
intensities for the devices were used to estimate the [BUP] in each
supernatant at each time increment (FIG. 15). This was accomplished
by using two separately generated calibration lines of [BUP] vs.
absorbance, including one at 270 nm (Table 12-3) and a second at
262 nm (Table 13-3). The two estimates of [BUP] were then averaged,
and were then used to assess the relative BUP release rates in
mg/ml/hour (Table 13-4 and FIG. 16), and the fraction of BUP
elution vs. time based on a theoretical maximum elution of
approximately 17 mg/ml for each of the delivery devices (Table 13-5
and FIG. 17).
[0356] Note that the raw absorption values from the supernatants of
the water-soaked devices were corrected for estimated background
contributions from non-BUP components that had become partially
dissolved over time. This was accomplished by subtracting the
absorbance values from the supernatant of the 14C-3E placebo from
those of the other devices at each respective wavelength as a
function of time. The data analyses were purposely limited to
include only those data time-points that were within the limits of
UV detection (i.e., below the saturation limit of the UV detector).
The upper time-value limits for each device are reported in Table
13-4. The detector saturation condition was reached more quickly
with devices that employed freely dispersed BUP powder (i.e., BUP
that was not encapsulated with PLGA). The maximum upper time-limit
in these cases was between approximately 12 and 48 hours. The
detector saturation condition was reached more slowly with delivery
systems that employed PLGA-encapsulated BUP microparticles (i.e.,
in devices made without the use of freely dispersed BUP). The
maximum upper time-limit before detector saturation in these cases
was approximately 96 hours (i.e., 4 days). Note that in some cases,
the estimates of [BUP] appeared to be slightly negative at short
soak times (e.g., at times of less than 8 hours for systems
formulated with PLGA-encapsulated BUP). This was an artifact of
over-correction from the 14C-3E placebo device, which appeared to
provide a slightly higher degree of short-time non-BUP component
dissolution than comparable devices that were formulated with
PLGA-encapsulated BUP.
[0357] FIG. 16 illustrates the relative rates of BUP elution
(mg/ml/hour) together with the data ranges used for establishing
the best linear fitting parameters. The initial slopes and data
ranges for the linear portions are reported in Table 13-4. Note
that for the purposes of these analyses, short-time negative
absorption values were omitted, a zero-time point was added (i.e.,
with absorption=0), and the best linear-fit lines were forced
through a zero-intercept. Data collection times included the
following time points (in hours): 1.5, 4, 8, 12, 24, 48, 72, 96,
120, and 192. However, only the on-scale data were represented
because beyond the upper time-point limit, the UV absorption values
moved off-scale due to detector saturation. FIG. 16 demonstrates
that for devices containing mixtures of freely dispersed BUP and
PLGA-encapsulated BUP, the relative BUP elution rate was
intermediate between rates for devices containing freely dispersed
BUP and for those containing PLGA-encapsulated BUP.
[0358] FIG. 17 mirrors the data plot presented in FIG. 16, but with
the [BUP] expressed in terms of the fraction of eluted
BUP=[BUP]/[BUP].sub.theoretical=[BUP]/17.14. This graph reveals
that the fastest releasing delivery system eluted 12% of its
theoretical [BUP] reservoir within about 12 to 24 hours, whereas
the slowest releasing systems released approximately 7 to 8% of
their theoretical [BUP] within approximately 4 days. Intermediate
devices (i.e., those containing mixtures of freely dispersed BUP
and PLGA-encapsulated BUP) had released 8-10% of their theoretical
[BUP] within approximately 2 days.
[0359] The formulations in each delivery device were formulated to
have similar percentages of total dispersed solids, components that
were not soluble in mineral oil but instead were dispersed within
the formulation matrix. Note that the total percentage of dispersed
solids is a factor that affects the rheological and compliance
characteristics of both the formulation and the delivery device.
These properties not only have an impact on the tactile
handleability of the delivery device during deployment, but they
also have an impact on the diffusion rates of fluids and active
ingredients as they diffuse across concentration gradients, both
into and out of the delivery device during its deployment and
during subsequent in vivo hydration. Generally, the higher the
percentage of dispersed solids, the higher the viscosity and the
lower the compliance. Of course, rheo-mechanical properties are
also affected by other factors, including for example, the particle
size distributions of the dispersed particulates, the total surface
to volume ratio of particulates within the formulation matrix and
within the delivery device, the weight and volume ratios of the
formulation to cellulose fibers in the delivery device, the number
and diameters of fibers that constitute a bundled-fiber strand, the
knit or weave density of the fibers and fiber bundles that
constitute a textile, and the surface wetting characteristics of
the fibers. Any one or combination of these factors can be
controlled and adjusted to achieve a broad range of rheo-mechanical
responses if so desired.
[0360] For the purposes of this example, dispersed solids were
calculated to include: 1) beeswax micro-crystallites dispersed in
mineral oil as a result of the melt-recrystallization process; 2)
bovine gelatin powder; 3) PLGA microspheres containing 20% BUP by
weight; 4) PLGA placebo microspheres; 5) BUP free base powder; 6)
citric acid powder; and 7) di-sodium citrate powder. By maintaining
similar levels of total dispersed solids, the resulting vehicles
were made to have qualitatively similar rheo-mechanical property
characteristics, including relative viscosity and compliance
characteristics as qualitatively judged by torque resistance during
spatula-mixing, and by compressibility during
textile-impregnation.
[0361] In one comparison, the relative BUP release rates were
compared among three types of formulations: 1) a formulation
containing PLGA-encapsulated BUP in a replicate of 14C-3A from
Example 12; 2) a formulation containing dispersed BUP free base
powder and placebo PLGA microspheres in a replicate of 14C-3B2 from
Example 12; and 3) a formulation containing a dispersed mixture of
both PLGA-encapsulated BUP and BUP free base (sample 14C-3C). The
14C-3C mixture resulted in a device with a relative BUP release
rate that was intermediate between that of 14C-3A with dispersed
PLGA-encapsulated BUP and that of 14C-3B2 with dispersed BUP free
base, thereby demonstrating one of the methods that can be used to
create formulations with controlled release characteristics (Table
13-4 and FIGS. 15, 16, and 17).
[0362] The PLGA-encapsulated BUP is reasoned to be slower to
diffuse and release because it encounters at least two diffusion
barriers, the first being the PLGA polymer itself, and the second
being the remainder of the formulation matrix. On the other hand,
dispersed BUP free base without PLGA encapsulation is thought to be
faster to diffuse and release because it encounters fewer diffusion
barriers. By mixing the two types of BUP at various weight ratios,
dispersed microspheres of encapsulated BUP mixed with dispersed
free base powder, it becomes possible to achieve a range of release
rates, any of which can be chosen to achieve a desired
control-release profile. Note that the optimum control-release
profile will depend upon the clinical needs of the end use
application.
[0363] It should be understood that this is only one method by
which one can achieve a controlled-release profile. Mixtures can be
augmented in other ways to include the use of other dispersed or
dissolved ingredients that can have an impact on release rates,
either alone or in combination with one another, or in combination
with the dispersed ingredients mentioned above, and at various
weight ratios. Other dispersed ingredients can include, for
example, BUP-HCl powder which is more water soluble than BUP,
PLGA-encapsulated BUP-HCl, PLGA-encapsulated mixtures of BUP free
base and BUP-HCl. Moreover, the same PLGA-encapsulated ingredients
can be comprised of larger or smaller PLGA particle size
distributions, or mixtures of different PLGA particle size
distributions.
[0364] Although there were differences in the dispersion
characteristics of the various types of solid ingredients due to
factors like particle size distribution and surface wetting
characteristics, it was still possible to adjust the ratios of
ingredients to achieve formulations with nearly equivalent levels
of total dispersed solids, while simultaneously maintaining
qualitatively similar compliance characteristics. Moreover, as
demonstrated by 14C-3C, higher net loadings of BUP were also
simultaneously achieved (Table 13-3). For the case of formulation
14C-3C, a higher BUP dosage was achieved by reducing the weight
percentage of bovine gelatin, and by then adding an equivalent
weight % of BUP free base powder in its place, so as to maintain a
similar equivalent percentage of total dispersed solids.
Importantly, this is an example of the type of formulation
flexibility that can allow for the creation of formulations with
the capacity to deliver higher maximum BUP dosages on a unit weight
basis than those that would otherwise be possible through the use
of PLGA microspheres alone, or more specifically with the 4.3
micron 20% BUP free base loaded PLGA microspheres as used in this
example. Moreover, the potential for higher maximum dosages in a
formulation with more hydrophobic ingredients will generally exceed
what is possible with the more hydrophilic formulation embodiments,
partly because the latter require water-dilution for plasticization
in order to render them as compliant and useable.
[0365] In a second group of comparisons, samples were formulated to
contain additional dispersed particulates of either citric acid or
di-sodium citrate. The purpose of employing these types of optional
pH modulators is to alter the relative acidity or basicity of the
local chemical environment during the hydration process. During in
vivo deployment, the formulation will absorb and mix with body
fluids from the tooth extraction socket (Example 9), and through a
process of diffusion, soluble components such as BUP-HCl, citric
acid, GLBG, SO, etc., will eventually leach out of the formula and
will become actively available to the surrounding tissues.
Modulators can serve multiple purposes, including, for example: 1)
to reduce or enhance the degree of BUP protonation affecting
solubility and chemical activity; 2) to neutralize acid hydrolysis
products (e.g., lactic acid that can form via hydrolysis of PLGA);
3) to reduce or enhance the degree of protonation of
gelatin-protein amines which can have an impact on rates of
gelation and property-build characteristics of gelatin during
hydration as demonstrated in earlier examples with citric acid; 4)
to form citrate salts that can exchange with and alter the
solubility or the chemical activity of conjugate acid-base pairs,
such as protonated BUP with Cl as its base-conjugate in exchange
with citrate as its base-conjugate; 5) to catalyze the hydrolysis
of PLGA, thereby enabling faster release rates of active
ingredients if so desired; and 6) to positively impact the tissue
healing process during end use, which is a known attribute of acids
like citric acid, ascorbic acid, and others. Importantly,
modulators that are insoluble in the liquid carrier, like mineral
oil, can be directly dispersed within the formulation. These
modulators can also be microencapsulated themselves within PLGA or
with other polymers and then can be dispersed in the formulation.
Of course, the purpose of microencapsulation would be to augment
their time-controlled availability to satisfy any of the
aforementioned purposes 1 through 6 as stated above.
[0366] It can be appreciated that when a modulator has an impact on
rheo-mechanical properties via reducing or enhancing the degree of
gelatin-protein amine protonation, it will consequently have an
impact on rates of diffusion and on the release rates of active
ingredients. Similarly, when a modulator has an impact on the
solubility of an active ingredient via formation of alternate
conjugate base pairs, or via direct protonation or de-protonation,
rates of diffusion and rates of release can be similarly
affected.
[0367] Examples of the use of modulators are represented by samples
14C-3E, 14C-3F, and 14C-3G. Each of these samples demonstrates the
formulation flexibility afforded by the more hydrophobic
formulation embodiment. Specifically, by using the more hydrophobic
formulation impregnated into a cellulose textile, five desirable
end use attributes were simultaneously and synergistically
demonstrated, including: 1) the ability to achieve specific
control-release profiles through the use of mixtures of dispersed
ingredients (e.g., PLGA-encapsulated BUP mixed with
non-encapsulated BUP); 2) the ability to achieve a wider range of
BUP dosage levels; 3) the ability to use mixtures to achieve higher
net BUP dosage levels than would otherwise be possible with
PLGA-encapsulated BUP alone; 4) the ability to achieve compliance
and tactile characteristics commensurate with those desired for the
end use application; and 5) the ability to achieve additional end
use functionality via the incorporation of dispersed pH modulators,
without negatively impacting the rheo-mechanical properties or the
efficacy of the device.
TABLE-US-00028 TABLE 13-1 Weight % compositions of hydrophobic
vehicles for use in preparing textile-impregnated devices.
Calculations also include the net weight % concentration of BUP in
each vehicle, the net PLGA polymer weight % (i.e., ~80% of the
weight of BUP loaded microspheres, and 100% of placebo
microspheres), and the total weight % of dispersed solids. Vehicle
Mixture 14C-3A 14C-3B2 14C-3E Composition Replicate Replicate
14C-3C 14C-3E 14C-3G 14C-3F Placebo Mineral Oil 23.03% 28.39%
23.03% 23.03% 23.03% 23.03% 28.12% Beeswax 4.61% 5.68% 4.61% 4.61%
4.61% 4.61% 5.62% Bovine Gelatin 13.85% 17.07% 9.24% 11.55% 6.94%
11.55% 10.57% 5 um PLGA Placebo 0% 39.09% 0% 0% 0% 0% 53.58%
microspheres (dispersed) 4.3 micron 20% BUP 58.51% 0% 58.51% 58.51%
58.51% 58.51% 0% free base loaded PLGA microspheres (dispersed) BUP
free base (directly 0% 9.77% 4.61% 0% 4.61% 0% 0% dispersed in
vehicle) citric acid (dispersed) 0% 0% 0% 2.30% 2.30% 0% 2.11%
di-sodium citrate 0% 0% 0% 0% 0% 2.30% 0% (dispersed) TOTAL 100.00%
.sup. 100% .sup. 100% .sup. 100% .sup. 100% .sup. 100% .sup. 100%
Total BUP in vehicle 11.70% 9.77% 16.31% 11.70% 16.31% 11.70% 0%
Total PLGA polymer in 46.81% 39.09% 46.81% 46.81% 46.81% 46.81%
53.58% Vehicle Total % dispersed solids 76.97% 71.61% 76.97% 76.97%
76.97% 76.97% 71.88% in vehicle
TABLE-US-00029 TABLE 13-2 Weight % compositions of hydrophobic
textile-impregnated devices. The vehicle compositions as reported
in Table 13-1 were impregnated into two orthogonally oriented SO
textiles. The calculations for compositions also include the
concentration of BUP per unit weight of device, and the effective
available BUP concentration for release during the water-soak
experiments. Note that when the devices were transferred to 11 ml
glass vials, a small amount of vehicle weight was lost. This loss
was taken into account to insure that the correct water to BUP
weight ratios were employed during the water-soak experiment (i.e.,
to achieve a BUP reservoir concentration of approximately 17 mg
BUP/ml water, which was the same concentration that was used during
the water-soak experiments in Example 12). 14C-3A 14C-3B2 14C-3E
Ingredient Replicate Replicate 14C-3C 14C-3E 14C-3G 14C-3F Placebo
Great Lakes Bovine 12.35% 15.36% 8.24% 10.35% 6.13% 10.28% 9.39%
Gelatin (GLBG) Mineral Oil (MO) 20.53% 25.54% 20.54% 20.63% 20.33%
20.51% 24.96% Beeswax (BW) 4.11% 5.11% 4.11% 4.13% 4.07% 4.10%
4.99% PLGA polymer (i.e., 41.73% 0% 41.74% 41.94% 41.31% 41.69% 0%
representing 80% of the weight of 4.3- micron microspheres loaded
with 20% by wt. BUP) Encapsulated BUP 10.43% 0% 10.43% 10.48%
10.33% 10.42% 0% (i.e., representing 20% by weight of the
4.3-micron microspheres loaded with 20% by wt. BUP) PLGA polymer
from 0% 35.16% 0% 0% 0% 0% 47.56% 5-micron placebo microspheres BUP
free base (non- 0% 8.79% 4.11% 0% 4.07% 0% 0% encapsulated,
directly added to the vehicle) Citric Acid 0% 0% 0% 2.06% 2.03% 0%
1.87% di-Sodium Citrate 0% 0% 0% 0% 0% 2.05% 0% SO textiles 10.84%
10.05% 10.83% 10.41% 11.74% 10.94% 11.24% mg BUP/g device 104 88
145 104 145 104 0 Target ratio of pH 2 58.33 58.33 58.33 58.33
58.33 58.33 Water/ water to BUP (w/w) in device water soak
experiment (w/w) ~14C-3A Weight of Device as 0.8134 0.8688 0.8483
0.8543 0.8064 0.8334 0.8222 made (g) Weight of Vehicle 0.7252
0.7815 0.7564 0.7654 0.7117 0.7422 0.7298 as made (g) Tarred Weight
(g) of 0.8037 0.8533 0.8409 0.8438 0.7953 0.8234 0.8105 device
added to 11 ml glass vial Weight of vehicle after 0.7155 0.7660
0.7490 0.7549 0.7006 0.7322 0.7181 transfer to vial (g) (containing
(containing (16.31% (11.70% (16.31% (11.70% (0% 11.70% 9.77% BUP)
BUP) BUP) BUP) BUP) BUP) BUP) Weight of pH 2 water 4.884 4.365
7.125 5.153 6.665 4.998 4.880 (g) added to 11 ml vial mg of
available BUP 17.14 17.14 17.14 17.14 17.14 17.14 0 per ml pH-2
water (i.e., [BUP].sub.theoretical)
TABLE-US-00030 TABLE 13-3 BUP calibration equation as obtained from
a linear best fit of absorption at 262 nm vs. BUP concentration
(mg/ml) in pH-2 water. This table provides the absorbance intensity
for BUP free base that was fully dissolved in pH-2 water over the
detectable range of [BUP], expressed in mg/ml. Note that the
absorbance values as reported below represent corrected values that
were obtained by subtracting a single-beam absorbance spectrum of
pH-2 water from the single-beam absorbance spectra of the BUP
samples. This calibration was used together with the calibration at
270 nm (Table 12-3) to estimate the [BUP] in supernatants from pH-2
water- soaked devices for the present example. [BUP] mg/ml Relative
Absorbance Intensity 2.2364 3.0275 1.2300 2.0130 0.72684 1.1347
0.22364 0.2393 0.022364 -0.0681 0.0022364 -0.1004 0.00022364
-0.0825 0 0 R.sup.2= 0.987 Slope= 1.4482 y-intercept= -0.0336
TABLE-US-00031 TABLE 13-4 Relative rates of BUP elution
(mg/ml/hour) established from the slopes of the linear portions of
each elution curve in FIG. 16. Note that for the purposes of these
analyses, short-time negative absorption values were omitted, a
zero-time point was added (i.e., with absorption = 0), and the best
linear fit lines were forced through a zero-intercept. This table
also includes the linear ranges that were used to obtain the best
linear fits, as well as the upper time limits that were used for
presentation of the data in FIG. 16. Data collection times included
the following time points (in hours): 1.5, 4, 8, 12, 24, 48, 72,
96, 120, and 192. For times above the upper-time limit, the UV
absorption values were off-scale due to detector saturation. Note
that for devices containing mixtures of freely dispersed BUP and
PLGA-encapsulated BUP, the measured rate (e.g., 14C-3C~0.06
mg/ml/hour) was observed to be in reasonable agreement with a
calculated rate that was based on weight fractions of freely
dispersed BUP and PLGA-encapsulated BUP multiplied by the rates
associated with devices that were formulated exclusively with
PLGA-encapsulated BUP, and exclusively with dispersed BUP (e.g.,
0.28 .times. (rate for 14C-3B2) + 0.72 .times. (rate for
14C-3A)~0.08 mg/ml/hour). Calculated Rate based on weight Relative
Rate fractions of Upper time limit for of BUP elution freely
dispersed data presentation in from FIG. 16 BUP and PLGA- Linear
best-fit FIG. 16 (i.e., UV Sample (mg/ml/hour) encapsulated BUP
time region detection on-scale) 14C-3A replicate 0.0137 NA 0 to 96
hours 96 hours with BUP encapsulated by PLGA 14C-3B2 replicate
0.2533 NA 0 to 8 hours 12 hours with freely dispersed BUP 14C-3C
0.0564 0.08 = 0.28(14C-3B2) + 0 to 24 hours 48 hours with a ~28/72
blend (w/w) 0.72(14C-3A) of BUP that was freely dispersed together
with BUP that was encapsulated with PLGA 14C-3E 0.0121 NA 0 to 96
hours 96 hours with BUP encapsulated by PLGA; and with dispersed
citric acid 14C-3G 0.0523 0.08 = 0.28(14C-3B2) + 0 to 24 hours 48
hours with a ~28/72 blend (w/w) 0.72(14C-3E) of BUP that was freely
dispersed together with BUP that was encapsulated with PLGA; and
with dispersed citric acid 14C-3F 0.0164 NA 0 to 72 hours 96 hours
with BUP encapsulated by PLGA; and with dispersed sodium
citrate
TABLE-US-00032 TABLE 13-5 The devices were allowed to continue
soaking in pH-2 water beyond the upper time limit that was reported
in Table 13-5, and an estimate of the fraction of BUP released at t
= 192 hours (8 days) was performed. At t = 192 hours, the
supernatants of all samples, including the 14C-3E placebo were
sampled, and were then subjected to a 10-fold dilution with pH-2
water. The dilution of the supernatants enabled the acquisition of
on-scale UV absorption spectra. The resulting absorbance values at
262 nm and at 270 nm were background corrected by using the 10-fold
diluted UV spectrum of the comparable 14C-3E placebo. The corrected
absorption values were then used to estimate the BUP concentrations
that had eluted into the closed systems at t = 8 days. The averages
of the values calculated from the 262 nm and 270 nm wavelengths
(using the calibration lines from Tables 12-3 and 13-3) are
presented below, together with the weight fractions that had eluted
after 8 days of soaking in pH-2 water. Estimated Estimated fraction
of [BUP] BUP released after released after t = 8 days = [BUP]/ t =
8 days [BUP].sub.theoretical = Sample (mg/ml) [BUP]/17.14 14C-3A
replicate 4.85 0.28 with BUP encapsulated by PLGA 14C-3B2 replicate
15.98 0.93 with freely dispersed BUP 14C-3C 7.65 0.45 with a ~28/72
blend (w/w) of BUP that was freely dispersed, together with BUP
encapsulated with PLGA 14C-3E 3.89 0.23 with BUP encapsulated by
PLGA, together with dispersed citric acid 14C-3G 8.22 0.48 with a
~28/72 blend (w/w) of BUP that was freely dispersed and BUP
encapsulated with PLGA, plus dispersed citric acid 14C-3F 9.12 0.53
with BUP encapsulated by PLGA, together with dispersed sodium
citrate
Example 14. Suspension Test for Choosing Liquid Components Suitable
for Use in Preparing Hydrophobic and Hydrophilic Formulations
[0368] As noted previously, hydrophobic formulations and delivery
devices can be desirable from the standpoint that they can be
formulated to yield dough-like materials with compliance
characteristics that are conducive to end use deployment, without
having to rely upon pre-deployment swelling and gelation of the
gelatin particulates. Thus, hydrophobic formulations and delivery
devices are ones whereby the gelatin particulates remain intact
during manufacture and during storage, and do not yield macroscopic
chain-entangled gelled networks until they become exposed to the
tooth extraction socket and its fluids after deployment, unless the
option of pre-deployment hydration is exercised.
[0369] It is important to note that each of the embodiments of the
formulation will eventually become hydrated with fluids from the
tooth extraction socket after deployment. This is predominantly due
to the presence of hygroscopic, water-absorbing network-forming
polymers like gelatin or to the presence of other water-absorbing
materials such as cellulose fibers. However, in order to render the
devices as compliant and conformable prior to their deployment, it
is desirable that they be properly formulated in advance of
deployment so that the clinician does not have to spend time
meticulously measuring and premixing materials before they can be
used. In other words, it is desirable to have a device that is
already a compliant solid without having to be premixed with fluids
like saline solutions or water.
[0370] In previous examples pertaining to the more hydrophilic
embodiment of the present formulation, water was used as a
plasticizer to pre-hydrate and to masticate blends of powdered
ingredients to yield compliant dough-like mixtures, including water
and bovine gelatin with PLGA-encapsulated BUP as described in
Example 12. In these cases, water was the primary liquid ingredient
in the formulation, and the mechanical integrity of the device was
achieved by virtue of gelation and network formation prior to the
deployment of the device. The compliance and conformability of
these formulations were controlled by the weight ratio of water to
gelatin with consideration also given to the total weight % solids
in the plasticized mixture. Importantly, water was used as a liquid
plasticizer for the gelatin polymer.
[0371] A plasticizer is generally a liquid (sometimes a solid) that
when blended with a polymer increases the fraction of free volume,
which in turn lowers the polymer glass transition temperature and
consequently the elastic modulus and increases the compliance.
Plasticizers are known to be at least partially miscible with the
polymers that they plasticize.
[0372] By contrast, in examples pertaining to the more hydrophobic
embodiment of the present formulation, oils with optional waxes
were used as liquid carriers to suspend hygroscopic,
water-absorbing network-forming polymers such as gelatin powders
together with other dispersed ingredients, including
PLGA-encapsulated BUP, free BUP, and citric acid, to name a few.
These devices achieved their pre-deployment conformability and
compliance characteristics not by plasticization of a polymeric
continuous phase, but instead by virtue of other interactive
factors that impact the rheological properties of suspensions,
including the ratio of hydrophobic liquid to wax, which controls
the viscosity of the liquid carrier and affects the viscosity of
the resulting vehicle, the particle size distributions of dispersed
ingredients, and the total percentage of dispersed solids in the
vehicle, to name a few. In these cases, the mechanical integrity of
the pre-deployed device was not achieved by virtue of gelling a
polymer with a plasticizer to yield a reinforcing polymer network,
but instead it was achieved by virtue of fiber reinforcement by
impregnating knitted or woven cellulose textiles, or non-woven
fibers with non-gelled suspensions to yield fiber-reinforced
composite-like structures.
[0373] Thus, one of the primary distinctions between the
hydrophilic and hydrophobic devices relates to pre-deployment
morphology. By design, a hydrophilic device is comprised of a
water-miscible hygroscopic polymer network that is homogenously
gelled and pre-plasticized with a polar, hydrogen bonding liquid
such as water, glycerin, honey, polyethylene glycols, polypropylene
glycols, etc.; while by contrast, the hydrophobic device contains
inter-dispersed fibrous components and suspended particulates of
water-miscible and hygroscopic network-forming polymers like
gelatin that have the latent potential to form gelled networks once
exposed to water (i.e., after deployment), but in their
pre-deployment state, they are made to persist as morphologically
discrete entities suspended within and wetted by a hydrophobic
vehicle. By design, these devices do not rely on gelatin
plasticization and network formation (gelation) to achieve their
pre-deployment properties. However, after deployment, they are
morphologically designed to accept water through diffusion, which
allows for post-deployment polymer network formation, analogous to
what occurs in the pre-deployment stage with a hydrophilic device.
At that point (i.e., after the deployment), the development of a
gelled polymer network from water-ingress can have the added
benefit of providing an additional mechanism of mechanical
reinforcement, augmenting that which is already provided by the
inter-dispersed cellulose fibers.
[0374] With these morphological considerations in mind, the
differences between a more hydrophilic and a more hydrophobic
device can be further reduced to another important
design-controlling distinction, namely, the nature of the liquid
component that is used in formulating the vehicle for the device.
Generally, a liquid that leads to pre-deployment gelation is best
suited and preferred for use in preparing the more hydrophilic
formulations. A liquid that does not lead to pre-deployment
gelation, at least little to no gelation for a period of time after
manufacture that coincides with the desired shelf-life of the
device prior to its deployment, is best suited and preferred for
use in preparing the more hydrophobic formulations. The delineation
between a liquid that leads to gelation and one that does not lead
to gelation can be defined by a suspension test as demonstrated in
the present example.
[0375] The miscibility of a liquid carrier with gelatin and hence
the propensity for gelation can be gauged with a simple suspension
test, where gelatin particulates are first blended with the liquid
at weight ratios sufficient to form pourable suspensions, for
example 2/1, 3/1, 4/1 or even higher weight ratios of liquid to
gelatin, including 10/1, 25/1 or more. The suspensions are then
qualitatively monitored as a function of time for physical changes,
such as the onset of gelation, by using any one of a variety of
possible qualitative or quantitative techniques. Note that other
liquid to gelatin ratios can also be employed, including ratios
that are intended for use in various end-applications. The ratio
that was used in the present example, 2/1 w/w liquid to GLBG, was
meant only to illustrate the phenomenon and to provide a general
rubric for making an educated choice pertaining to liquid
carrier.
[0376] Monitoring times of suspensions can include various time
points after the suspensions are mixed, including for example, 5
minutes after mixing, 0.5 hours after mixing, 1 hour after mixing,
5 hours after mixing, 24 hours after mixing, 48 hours after mixing,
1 week after mixing, 1 month after mixing, 6 months after mixing, 1
year after mixing, and even 2 to 5 years after mixing.
[0377] Qualitative techniques for monitoring suspensions for
time-dependent changes that pertain to gelation or the lack thereof
can include, for example: 1) monitoring the suspensions for
relative time-dependent changes in viscosity by means of
hand-stirring the suspensions with a spatula (spatula test-1); 2)
determining whether the suspensions can still be poured from their
containers after various periods of aging (pour test); 3) shaking
the suspensions by hand to determine if they remain as liquid
dispersions after various periods of aging (shake test); 4) using
an optical microscope to determine if discrete particulates of
gelatin remain intact and suspended within the liquid over time
(microscope test); or 5) qualitatively viewing the elastic recovery
response of the suspension when it is perturbed by hand using
either a spatula or a similar object to see if a plasticized and
gelled polymer network begins to develop, as evidenced by being
able to lift the plasticized polymer from its container with a
spatula (spatula test 2). Of course, quantitative measurements can
also be employed if so desired (e.g., Brookfield viscosity,
parallel plate dynamic mechanical techniques, etc.). The tests
should be performed at a temperature that is above the melting
point of the liquid so that the resulting suspension is initially
one that is characterized as having particulates dispersed in a
liquid as opposed to particulates dispersed in a solid.
[0378] Using any one of these qualitative techniques, candidate
liquids for a more hydrophobic device are those that when mixed
with gelatin-particulates form suspensions that exhibit one or more
of the following responses within a preferred monitoring time
window: 1) minimal to no change in relative viscosity (spatula
test-1); 2) retention of pourability (pour test); 3) retention of
liquid dispersion/suspension state characteristics (shake test); 4)
maximum retention of discrete gelatin particulates within the
liquid continuous phase (microscope test); and 5) minimal to no
elastic recovery (spatula test-2). In general, if there are no
signs of gelation as gauged by one or more of these responses
within a preferred monitoring time window, then these liquids are
considered to be candidates for use in preparing a more hydrophobic
device. Mineral oil, caprylic triglyceride, isopropyl palmitate,
and coconut oil are such liquids as illustrated by the results in
Table 14-3.
[0379] Conversely, if there are signs of gelation, including for
example, any one or more of the following responses before the end
of the monitoring time window: 1) development of elastic recovery;
2) visualization of permanent coalescence of gelatin particulates;
or 3) solid network formation with loss of pourability, then these
liquids are by definition excluded as candidates for use in more
hydrophobic formulations, and are instead considered as candidate
liquids for use in preparing more hydrophilic formulations. Thus,
liquids that are observed to lead to gelation of gelatin within the
preferred time monitoring window are considered to be good
candidates for use in preparing a more hydrophilic formulation.
Glycerin and water are such liquids as shown in Table 14-3.
[0380] For the purposes of this example, the preferred monitoring
time windows for the suspension test are 0-24 hours and 0-48 hours.
For preparing a more hydrophobic formulation with optimal storage
stability, the preferred monitoring time is more preferably 0-3
months or 0-6 months, and even more preferably, 0-12 months.
[0381] If the particulates of a water-miscible and hygroscopic
network-forming polymer do not gel for at least a time period of 24
hours after being separately suspended within one or more
hydrophobic components, or more preferably for at least a time
period of less than 3 to 12 months after being separately suspended
within one or more hydrophobic components, then the one or more
hydrophobic components are deemed suitable for preparation of a
hydrophobic sustained release system or formulation. Conversely, if
the particulates of a water-miscible and hygroscopic
network-forming polymer form a gel within a time period of 24 hours
after being separately suspended within one or more of the
components, then the one or more components are deemed suitable for
preparation of a hydrophilic sustained release system or
formulation. Importantly, if the hydrophilic sustained release
system is intended to be reinforced with a fibrous member such as a
cellulose textile, then it is preferable that the particulates of
the water-miscible and hygroscopic network-forming polymer do not
gel for at least a time period of 2 hours, and more preferably for
a time period of at least 4 to 8 hours after being separately
suspended within one or more hydrophilic components. In this way,
the delivery system can be more readily manufactured within a
manageable work-time window, wherein the viscosity of the
formulation remains sufficiently low enough to facilitate
impregnation of the reinforcing member.
[0382] For desirable embodiments where particulates of a
water-miscible and hygroscopic network-forming polymer do gel in at
least a time period of 24 hours after being separately suspended
within one or more hydrophobic components, two component,
mix-on-demand such as two part syringes would be considered to
realize these desirable embodiments.
[0383] Note that similar tests can be employed to test the
miscibility of carrier liquids with other dispersed ingredients,
including for example, microspheres of PLGA-encapsulated BUP,
freely dispersed BUP, freely dispersed BUP-HCl, citric acid,
ascorbic acid, citrates, etc. The liquid carrier can also be
modified in advance of the test via incorporation of optional waxes
or surfactants if so desired.
[0384] Importantly, the suspension test is not necessarily limited
to gelatin protein. Instead, it can be used to test the suitability
of a liquid for use in preparing hydrophobic or hydrophilic
formulations wherein the formulation comprises other, alternative
water-miscible and hygroscopic network-forming polymer components
besides gelatin. Thus, in its most general sense, it is intended to
test the suitability of a liquid for preparing hydrophilic or
hydrophobic formulations whereby the formulation contains a
hygroscopic, water-absorbing network-forming polymer component,
such as a protein polymer like gelatin, or other alternative
hygroscopic network-forming polymer components, including natural
gums from a variety of plant sources, such as tree exudates of
which arabic, ghatti, karaya, and tragacanth are examples, seaweed
colloids of which agar, Irish moss, carrageenin, and alginates are
examples, seed extracts of which locust bean, locust kernel, and
quince seed gums are examples, manufactured and modified dextrins,
British gums, and water-dispersible or soluble derivatives of
cellulose to name a few. A more thorough account of these and
similar materials can be found in The Water Soluble Gums, C. L.
Mantell, Reinhold Publishing Corporation, New York, 1947. Thus,
independent of which hygroscopic, water-absorbing network-forming
polymer is chosen, particulates of the polymer are dispersed in a
test-liquid, and the suspension test is conducted using the same
procedures as those outlined for gelatin in the present
example.
[0385] In this example, suspension tests were conducted using
candidate liquid carriers as described in Table 14-1. 0.50 g
aliquots of GLBG with general information provided in Table 14-2
were weighed into 11 ml glass vials with lids. Next, a 1 g aliquot
of a candidate test liquid was weighed into an individual vial
containing the GLBG to achieve a 2/1 liquid/GLBG weight ratio. A
spatula was used to stir the ingredients to create a suspension.
The suspensions were then allowed to set under static conditions
and were qualitatively monitored as a function of time. Results at
t=5 minutes after mixing, t=0.5 hours after mixing, t=5 hours after
mixing, t=24 hours after mixing, and at t=48 hours after mixing
were reported. The tests were conducted at 20 degrees C. with one
exception, one of the tests was conducted at 27 degrees C. to
ensure that the carrier was above its melt point and in its liquid
state. For cases where sedimentation was observed to occur, which
happened over time with liquids that did not lead to gelation, the
spatula and shake tests were used to facilitate redispersion of the
gelatin particulates so that pourability could also be evaluated.
Results are provided in Table 14-3.
TABLE-US-00033 TABLE 14-1 Liquids used for suspension tests. Liquid
Distilled water Glycerin; USP grade, 99.9% anhydrous; Rite-Aid; CAS
# 56-81-5 caprylic triglyceride; Croda, Inc.; CAS # 65381-09-1 (see
Example 10) isopropyl palmitate; Sigma-Aldrich; CAS # 142-91-6 (see
Example 10 coconut oil (virgin); Nutiva; cold-pressed unrefined;
UPC 692752200052; CAS# 8001-31-8; melt point 76 deg. F. mineral
oil; Aldrich; CAS 8042-47-5 (see Example 9)
TABLE-US-00034 TABLE 14-2 Analytical data and specifications for
the Great Lakes brand of bovine gelatin (GLBG) that was used in the
suspension tests. Manufacture: Bovine gelatin powder, Great Lakes
Gelatin Company, Grayslake, IL, type B (bovine, alkali process),
unflavored Kosher beef hide, 88-92% protein, Kosher, Gluten Free,
US Pharmacopeia consumer grade General Analysis: PROTEIN 88-92%
Bloom 225 g Viscosity mp 34-40 pH 4.1-5.5 Moisture <12% Ash
<2% Sodium 100 mg/100 g Carbohydrates 0% Fat 0% Heavy Metals
<0.005% Bacteria Test USP/NF Calories per ounce 103.0 Maximum
Amino Acid Content: Alanine 11.0%/1,210 mg Arginine 9.3%/1,023 mg
Aspartic Acid 6.7%/737 mg Cystine 0.1%/11 mg Glutamic Acid
11.4%/1,254 mg Glycine 29.0%/3,190 mg Histidine 1.0%/110 mg
Hydroxyproline 14.5%/1,595 mg Hydroxylysine 1.2%/132 mg Isoleucine
1.8%/198 mg Leucine 3.4%/374 mg Lysine 4.6%/506 mg Methionine
1.0%/110 mg Phenylalanine 2.6%/286 mg Proline 17.6%/1,936 mg Serine
3.8%/418 mg Threonine 2.2%/242 mg Tryptophane 0.0%/0 mg Tyrosine
1.0%/110 mg Valine 3.3%/363 mg
TABLE-US-00035 TABLE 14-3 Suspension test results. Each suspension
existed as a liquid dispersion at t = 0. Results at either t = 5
minutes after mixing, t = 0.5 hours after mixing, t = 5 hours after
mixing, t = 24 hours after mixing, or t = 48 hours after mixing are
reported. Suitable for Suitable for use in a use in a T of Test
Spatula Pour test & Spatula hydrophobic hydrophilic Liquid
(deg. C.) test-1 Shake test test-2 device device Water 20 High
viscosity at Neither pourable nor Elastic network at No Yes t = 5
minutes shakable at t = 5 t = 5 minutes minutes Glycerin 20 No
change at 5 min.; Pourable and shakable Elastic network at No Yes
waxy dispersion at at 5 minutes but not 24 hours 5 hours at 5 hours
caprylic 20 No change, a liquid No change, pourable No elastic
network Yes No triglyceride dispersion from 0-48 and shakable
dispersion formation from 0-48 hours from 0-48 hours hours
isopropyl 20 No change, a liquid No change, pourable No elastic
network Yes No palmitate dispersion from 0-48 and shakable
dispersion formation from 0-48 hours from 0-48 hours hours coconut
oil 27 No change, a liquid No change, pourable No elastic network
Yes No dispersion from 0-48 and shakable dispersion formation from
0-48 hours from 0-48 hours hours mineral oil 20 No change, a liquid
No change, pourable No elastic network Yes No dispersion from 0-48
and shakable dispersion formation from 0-48 hours from 0-48 hours
hours
[0386] In some circumstances, the degree of hydrophilicity and
hydrophobicity of a liquid can also be gauged by parameters that
pertain to molecular-level properties such as polarity (e.g.,
dipole moment forces from permanent dipoles), dispersion forces
(e.g., non-permanent dipoles or van der Waals forces), and hydrogen
bonding forces. Indices such as the Hildebrand Solubility Parameter
(HSP) or Hansen Solubility Parameter (HAN) of liquids and polymers
(J. Brandrup and E. H. Immergut, Polymer Handbook, Third Edition,
John Wiley & Sons, New York, 1989, pp. 519-559), as well as Hoy
solubility parameters (HOY), have been developed in attempts to
better quantify what is meant by "hydrophilicity" and
"hydrophobicity." Hoy solubility parameters (HOY), like Hansen
Solubility parameters (HAN) are based on chemical group methods of
calculating energetic contributions from dispersion forces, polar
forces, and hydrogen bonding forces. These contributions are summed
to yield the total solubility parameter by taking the square root
of the sum of the squares. Generally, although the estimation
methods differ for the HAN and HOY terms, the sums of the
contributions from HAN and HOY parameters produce similar total
solubility parameter estimates, which are also considered to be
equivalent to HSP values (i.e.,
HSP.about.HAN.sub.total.about.HOY.sub.total).
[0387] It is generally understood by those skilled in the art that
polymers and liquids tend to be more miscible when their solubility
parameters are similar in magnitude to one another. Conversely,
polymer/solvent pairs become less miscible as their solubility
parameters diverge from one another.
[0388] Various solubility parameter values as reported in the
literature for components like those found in the present
formulations are provided in Table 14-4.
[0389] For the purposes of the present description, the most
hydrophobic liquids can be defined as those with either a small or
no permanent dipole moment, and with a low capacity to participate
in hydrogen bonding. These types of liquids have been observed to
be the least compatible with highly polar and water-soluble
protein-based polymers like gelatin, which explains why the gelatin
particulates remain dispersed and stable over time when suspended
(i.e., not gelled) in formulations comprising such liquid carriers.
These types of liquids would also be expected to have limited
compatibility with other polar molecules such as water and BUP-HCl,
thus rendering them as relative deterrents to both molecular-level
and macro-level diffusion during the end use application as has
been illustrated in Example 12. This behavior renders such liquids
as useful levers in quests aimed at achieving specific control over
time-release profiles. An example of an extreme version of this
type of liquid is represented by a paraffinic hydrocarbon like
mineral oil.
[0390] On the other side of the spectrum, liquids with permanent
dipoles and with higher capacities for hydrogen bonding can be
classified as being less hydrophobic and more hydrophilic. In the
present description, this type of liquid is represented by water in
one extreme (HSP=approximately 48 MPa.sup.1/2). These types of
liquids are highly compatible with hygroscopic polymers like
gelatin, which explains why the dispersed gelatin particulates do
not persist in formulations containing water, but instead become
swollen through diffusion and plasticization, leading to the
coalescence of the particulates through polymer chain entanglement
and leading ultimately to gelation and to solid network formation
prior to deployment of the device.
[0391] Note that for the case of a more hydrophobic formulation
that is prepared with hydrophobic components like oils or waxes,
the more hygroscopic components like gelatin particles and
cellulose fibers remain discrete and intact prior to hydration,
either as dispersed, non-gelled particulates, or as intermeshed
fibrous entities. In these cases, the oils and waxes that
constitute the continuous phase of the formulation serve to
facilitate the dispersion of other ingredients like gelatin, PLGA
microparticles, BUP, and citric acid. Note that optional
surfactants can also be added to assist in stabilizing such
dispersions.
[0392] In a pre-deployment morphological state, the mechanical
integrity of the more hydrophobic formulation is predominantly
derived from its reinforcement with cellulose fibers. Importantly,
the morphology of the hydrophobic formulation has been designed to
adsorb polar liquids like water as demonstrated in Examples 5 and
7. Thus, when a polar liquid, such as water, glycerin, polyethylene
glycol, mixtures thereof, or fluids from the tooth extraction
socket, etc., is intermixed with a more hydrophobic formulation,
the morphology of the formulation and of the delivery device
accommodate the adsorption of the polar liquid without producing
the side effect of macroscopic phase separation of other
components. This behavior is consistent with a morphological change
that occurs when polar liquids are mixed with the device, whereby
the more hygroscopic components like gelatin or cellulose begin to
absorb the polar liquid becoming plasticized, and then begin to
coalesce into a gelled network matrix such that the new continuous
phase contains the gelled network matrix (i.e., polar
liquid+gelatin+cellulose), inter-dispersed together with the
hydrophobic components, the oils and waxes that previously
constituted the continuous phase prior to hydration. At this stage,
other dispersed ingredients like PLGA, BUP, BUP-HCl, citric acid,
etc., that were previously dispersed in the oil-based continuous
phase, either remain dispersed within the oil-phase components that
themselves become inter-dispersed within the gelled matrix, or they
become directly dissolved in the water that diffuses into the
newly-formed continuous phase of the gelled matrix). Importantly,
the plasticization, the chain-entanglement, the ensuing gelation,
and the ultimate network formation that accompanies this adsorption
process are desirable attributes for the more hydrophobic
formulation. Most importantly, and by design, this morphological
change is made to occur in vivo and does not have to occur during
the pre-deployment stage or during the storage period for the
formulation.
[0393] The latent capacity for a hydrophobic device to adsorb a
polar H-bonding liquid like water is not only a desirable and
surprising attribute that arises from the synergistic interactions
among the component ingredients of the formulation, it is a
measurable attribute that can be used to specify a distinguishing
characteristic of a more hydrophobic formulation. Namely, a more
hydrophobic device is one that after being mixed via physical
mastication with water at a minimum ratio of water to device=0.2/1
w/w, or more preferably 0.33/1 w/w, or even more preferably 0.44/1
w/w or higher, does not exhibit macroscopic phase separation under
static conditions for a period of at least 1 hour, and preferably
for 2 or more hours, and more preferably for 24 hours or more. It
further retains the added water for said period of time under
static conditions without exhibiting visual indications of macro
phase separation of water or other components. Indeed, this
behavior was exemplified by fibrous textile-reinforced hydrophobic
delivery devices that were demonstrated in Examples 5 and 7.
[0394] As stated previously, if the end-product objective is to
minimize active-ingredient dilution in the formulation while
simultaneously achieving mechanical compliance characteristics that
are desirable for deployment, then gelation of gelatin or other
macromolecular hygroscopic components would be most desirable if it
were made to occur after deployment of the formulation and not
before. Thus, the formulation of a more hydrophobic formulation
with a hydrophobic liquid like mineral oil or others as shown in
Table 14-3 represents an approach towards achieving this
objective.
[0395] On the other hand, when compared to hydrophobic liquids like
mineral oil, hydrophilic liquids like water and glycerin are more
compatible and more miscible with polar molecules like BUP-HCl, a
fact which is consistent with the observation of faster diffusion
rates exhibited by formulations that are pre-plasticized with water
as opposed to those prepared with mineral oil as the liquid vehicle
carrier as in Example 12. Hence, if the end-product objective is to
maximize the release rates of water-soluble active-ingredients
while simultaneously achieving mechanical compliance
characteristics that are desirable for deployment, then
pre-gelation of gelatin or other hygroscopic components with
hydrophilic liquids like water and glycerin could be a desirable
approach wherein gelation is made to occur before deployment of the
device. Thus, the formulation of a more hydrophilic formulation
represents a method of approach towards achieving this objective,
but only if the resulting dilution of active ingredients can be
tolerated in the end use application.
[0396] Again, in the absence of gelation, the more hydrophobic
formulas achieve their initial mechanical cohesive integrity
through a mechanism that is independent of gelled network
formation. Specifically, if the formulation is formulated to have
the compliance characteristics of a cream, it can then be used to
disperse active ingredients, and it can then be impregnated into a
fibrous textile which serves as a reinforcing scaffold for the
formulation before its deployment. The reinforced delivery device
is therefore made to have cohesive integrity and compliance which
renders it as sufficiently acceptable for use by the clinician
during its deployment. It is only later, after deployment, that the
gelatin particulates dispersed within the formulation and cellulose
fibers begin to swell with liquids from the tooth extraction
socket, leading to their chain entanglement and ultimately to their
network formation and to an accompanying change in morphology. The
gelled network then becomes a type of reinforcing scaffold for the
device in vivo, serving to enhance its cohesive strength which
enhances its mechanical integrity after deployment and not
before.
[0397] Other liquids besides mineral oil, such as caprylic
triglyceride and isopropyl palmitate as demonstrated in Example 10,
are more polar than mineral oil, and they have at least some
capacity for hydrogen bonding. However, their polarity and
H-bonding characteristics are insufficient to cause gelation of the
gelatin particulates that are suspended within them. Thus, although
these types of liquids have permanent dipoles and therefore have
some capacity for hydrogen bonding, they are poor plasticizers for
gelatin. For the purposes of the present description, formulations
comprised of such liquids are also classified as more hydrophobic
formulations and delivery devices. These more hydrophobic
formulations and the delivery devices containing them have a
distinguishing attribute in common, the liquid carriers that serve
to suspend and bind the ingredients within the vehicle do not
promote the gelation of the gelatin particulates, and they are
either immiscible with gelatin or have limited miscibility under
ambient conditions. Consequently, macromolecular chain entanglement
and gelation do not occur when the particulates are suspended in
such liquids.
[0398] Liquids that are deemed as being suitable for use in a more
hydrophobic formulation via the suspension test can also perform
other functions when included in the formulation. For example, the
HAN of isopropyl palmitate is reported as 15.3 MPa.sup.1/2.
Although these types of liquids are recognized as being more polar
than mineral oil, for the purposes of the present description they
are still classified as being relatively hydrophobic in that they
do not diffuse and swell gelatin particulates in the way that water
does. Instead, the gelatin protein particulates persist in such
formulations until they are subjected to hydration during end use.
Nevertheless, the permanent dipole moments of these liquids would
be anticipated to render them as more amenable to facilitating
molecular-scale diffusion of small polar molecules than would
mineral oil. Thus, liquids of these types can be useful to modulate
diffusion rates of active ingredients, thereby providing an
additional lever to achieve intermediate controlled-release time
profiles. In addition, hydrophobic liquids with higher polarity
than mineral oil can also serve the secondary purpose of lowering
the Tg of PLGA via plasticization. This would result in a faster
rate of diffusion of encapsulated ingredients because a lower Tg
will equate to a higher fraction of free volume, which in turn
would translate to lower potential energy barriers for diffusion of
small molecules across the PLGA polymer gradient from within the
PLGA particle and into the binder matrix.
TABLE-US-00036 TABLE 14-4 Hildebrand Solubility Parameters (HSP),
Hansen Solubility Parameters (HAN), and Hoy Solubility Parameters
(HOY), as reported or estimated from J. Brandrup and E. H.
Immergut, Polymer Handbook, Third Edition, John Wiley & Sons,
New York, 1989, pp. 519-559; or as referenced from other footnoted
sources. Note that the total solubility parameter for the purposes
of the present invention is taken as any one of the following
values: HSP~HAN~HOY. HAN.sub.total HAN .delta..sub.Dispersion HAN
.delta..sub.Polar HAN .delta..sub.H-bonding HSP (or HOY.sub.total
if so (or HOY if so (or HOY if so (or HOY if so Material
MPa.sup.1/2 noted) MPa.sup.1/2 noted) MPa.sup.1/2 noted)
MPa.sup.1/2 noted) MPa.sup.1/2 Mineral oil.sup.c 15-18 (estimated)
16-18 (estimated) 16-18 (estimated) 0 (estimated) 0 (estimated)
isopropyl palmitate -- 15.3 14.3 3.9 3.7 Caprylic
triglyceride.sup.a -- 17.0 16.2 3.4 4 Glycerol 33.8 36.2 17.4 12.1
29.3 Water 47.9 47.9 15.5 16.0 42.4 water.sup.d -- 48.0 12.2 22.8
20.4 Coconut Oil.sup.a -- 16.6 16.2 2.5 2.8 PLGA (lactide/ -- 21.7
17.4 7.6 10.5 glycolide = 100/0).sup.b PLGA (lactide/ -- 21.7 17.4
8.3 9.9 glycolide = 85/15).sup.b PLGA (lactide/ -- 21.7 17.4 8.3
9.9 glycolide = 75/25).sup.b PLGA (lactide/ -- 22.3 17.4 9.1 10.5
glycolide = 50/50).sup.b Denatured Dry -- 22.5 11.7 12.1 14.8
Collagen (gelatin).sup.d Denatured Wet -- 30.1 11.8 15.3 22.5
Collagen (gelatin).sup.d .sup.aAnaid De La Pena-Gil, Jorge F.
Toro-Vazquez, and Michael A. Rogers, Food Biophysics, Springer
Science+Business Media, New York 2016. .sup.bSchenderlein, S.,
Luck, M., Muller, B. W., International Journal of Pharmaceutics 286
(2004) 19-26. .sup.cestimated from ranges attributed to other long
chain hydrocarbons as reported in J. Brandrup and E. H. Immergut,
Polymer Handbook, Third Edition, John Wiley & Sons, New York,
1989, pp. 519-559. .sup.dHoy solubility parameters as reported by
Pashley, David H., et al., American Journal of Dentistry, 20 (1),
2007, p. 9.
Example 15. Preparation of a Fibrous Reinforced Delivery Device
with Glycerin as the Liquid Component
[0399] There are occasions when the use of a formulation comprising
a hydrophilic liquid would be desirable for end use. For example, a
formulation that is pre-mixed with water can be useful in achieving
relatively fast time-release profiles of water-soluble ingredients
as demonstrated in Example 12. The present description provides for
creating a formulation that is first premixed and pre-plasticized
with water, glycerin, polyethylene glycols, other polyhydric
alcohols, or mixtures thereof. These types of formulations are
analogous to the more hydrophobic formulations, but they are made
with a polar H-bonding liquid as the primary liquid ingredient
instead of oils and waxes, and they are designed to gel prior to
deployment instead of afterwards. Thus, as long as they are
shelf-stable, these types of formulations can be used for
controlled-release delivery on their own without fiber
reinforcement. However, they can also be optionally reinforced with
a fibrous cellulose hemostat to form a composite structure. The
purpose of this example is to demonstrate this aspect of the
formulation.
[0400] As noted by Jaymin C. Shah and Manoj Maniar in Journal of
Controlled Release, 23 (1993) 261-270, control release of active
ingredients like BUP from polymeric matrices, such as biodegradable
polyanhydride polymers, can occur via diffusion, dissolution or
erosion of the polymer. The authors note that erosion or diffusion
processes are generally assumed to control the rate of drug
release. Hence, if the drug and its conjugate salt have low water
solubility, then it is anticipated that the dissolution rate of the
drug could have significant effect on the release-kinetics of the
drug.
[0401] It should also be realized that diffusion and erosion are
interactive processes, and that diffusion involves not just the
egress of active ingredients from a delivery device, but ingress of
water and fluids from the chemical environment where the device is
deployed. As fluids diffuse into the device via both macro and
molecular-level pathways, the matrix polymer can become more
susceptible to erosion, either through dissolution of volume
elements from the exposed surfaces of the delivery device, from the
macro separation of particulates near the surfaces of the device,
or through a combination of the two.
[0402] As noted earlier, one advantage of using fibrous
reinforcement for a delivery device is that it can improve the
cohesive integrity of the device, and thereby render it to be more
erosion resistant. When a delivery device erodes during end use,
internal cohesive failures of the matrix can cause particulates of
the device to become macroscopically separated from the original
structure. During end use, fluids can permeate into the matrix
phase of the device through a combination of macroscopic and
microscopic diffusion mechanisms. Macroscopic diffusion can occur
through permeable boundaries that are present from defects like
void elements arising from entrapped air between partially bonded
matrix polymer particulates, such as gelatin particulates, or from
matrix polymer that is partially delaminated from the surfaces of
weakly bonded elements or components that are dispersed within the
matrix.
[0403] If the matrix contains a polymer that is hygroscopic, as it
is in a more hydrophilic formulation, molecular level diffusion of
hydrous liquids can occur along every frontal boundary that becomes
available to the fluid. When the fluid macroscopically diffuses
into the matrix along a frontal boundary, it also can begin to
permeate into the matrix polymer through a process of
molecular-level diffusion. As a volume element of a matrix polymer
begins to expand from the ingress of lower molecular weight fluids,
it can become plasticized by the fluid, leading to an increase in
the fraction of free volume within the matrix polymer phase and to
a subsequent further increase in the rate of molecular level
diffusion, both into and out of the matrix polymer network.
[0404] An increase in free volume at the molecular level also leads
to a number of additional physical changes in the matrix polymer
phase, including a decrease in the glass transition temperature, an
accompanying decrease in modulus, a decrease in ultimate stress to
failure resulting in lower strength, and to an accompanying
acceleration in the rate of molecular level diffusion of molecules
both into and out of the matrix polymer phase. The macro volume
expansion of the liquid-occupied volume element, that is the
polymer volume element that has become diffusion-permeated and
plasticized by fluids, leads to the development of localized
stresses that tend to accumulate at weak boundaries, such as at
frontal boundaries that separate swollen volume elements from other
volume elements that have not yet been permeated and are not yet
swollen. Defects sites near these boundary regions become
particularly susceptible to localized stress-induced tensile and
shear types of failures. The ensuing number of internal cohesive
failure events can begin to increase and even to accelerate from
excessive strains at weak junctures at cell walls of macroscopic
voids, at the interfaces of weakly bonded particulates, etc. The
cycle continues as more macroscopic pathways develop for the
macroscopic ingress of even more fluids, leading to a further
increase in the number of pathways for molecular level diffusion,
which then leads to an increase in the number of swollen volume
elements, which then leads to the further development of more
localized stresses. Hence, the cascade continues, culminating in an
acceleration in the rate of occurrence of ultimate failure
events.
[0405] The interconnected processes of erosion and diffusion can
also affect the efficacy of a delivery device. Clearly, as erosion
occurs, the total amount of surface area simultaneously increases.
This will affect one of the primary functions of the device--to
achieve and maintain a specific time-controlled release profile of
one or more active ingredients during end use. An increase in the
total surface area from erosion leads to an acceleration of
molecular-scale diffusion of active ingredients across the growing
number of concentration gradients that are provided by the growing
number of interfacial boundaries. This process will not only impact
the molecular level diffusion rates through the matrix polymer, it
can impact the molecular level diffusion rates through other types
of secondary diffusion barriers that have been purposely put into
place, the diffusion barrier created by a PLGA polymer which serves
to impede the molecular-level diffusion rate of its encapsulated
active ingredients like BUP or BUP-HCl.
[0406] Any process that leads to an increase in free volume of a
polymer will subsequently lead to an increase in the number of
molecular pathways that are available for molecular level
diffusion. Importantly, diffusion of small molecules will occur
across passive boundaries where a concentration gradient is in
existence (i.e., Fickian diffusion). Aside from relative polarity
considerations, the rate of diffusion depends on the fraction of
free volume within the materials on both sides of the frontal
boundary, as well as the relative concentration of the diffusing
species on both sides. Thus, as fluids begin to have access to the
surfaces of PLGA particles within the delivery device, they can
permeate the surfaces of the particles and thereby increase free
volume, and then increase the rate of diffusion of small molecules
that are encapsulated and contained within them. To add even more
complexity to this scenario, if the fluid contains water, PLGA can
hydrolyze. The hydrolysis process leads to a decrease in molecular
weight, to the production of more chain ends, and thus to a further
increase in free volume which further enhances the rate of
diffusion. A gelatin matrix polymer with polypeptide sequences will
also be susceptible to the same type of hydrolysis-initiated
acceleration of free volume. Thus, each molecular level diffusion
barrier that is purposely set in place to control the release of
drugs and the like can become altered and affected by a cascade of
macroscopic and molecular-level events. These events will
collectively affect the global time release profile of the
device.
[0407] It is understood that, when harnessed for the purpose of
achieving specific control-release profiles over sustained periods
of time, these mechanisms can be useful. On the other hand, if
these processes occur too quickly, it may become difficult if not
impossible to achieve longer-term sustained release. As shown in
Example 12, this is most particularly the case for a more
hydrophilic delivery device.
[0408] Importantly, composite structures can be used to reduce the
rate of occurrence of internal cohesive failure events of the types
described above. In a composite-like structure, the matrix can be
reinforced with fibers or with particulates, which serve as
scaffolds that can help to hold a mechanically weaker matrix phase
in place by reducing the probability of crack growth and
propagation along any one single boundary via distributing stresses
from swelling over larger volume elements and hence over multiple
boundaries within the structure, thereby reducing the magnitudes of
localized stresses and strains, and hence reducing the number and
frequency of catastrophic failure events. Lower levels of localized
stresses will translate to lower localized strains, which in turn,
depending on the geometric structure of the defect site, can lead
to sustained mechanical and cohesive integrity of the delivery
device over longer periods of time.
[0409] The more hydrophobic formulations lend themselves well to
the creation of fiber-reinforced composites primarily because, by
design, the formulations that are used to impregnate the fibers are
not pre-gelled into macro polymeric networks. Instead, these
formulations, with their hydrophobic liquid carriers, remain
compliant and moldable for long periods of time. The gelatin
particulates suspended therein do not begin to gel and swell until
they are exposed to fluids within the tooth extraction socket. Even
then, the rate of water ingress is diminished owing to the
hydrophobic nature of the formulation. All of this translates to an
extended work-time for accomplishing the manufacturing steps that
are required to make a composite device, including the time needed
to complete multiple process steps, such as mixing, metering,
impregnating, conveying, cutting, and packaging.
[0410] On the other hand, the creation of a composite reinforced
delivery device that is more hydrophilic poses a different set of
challenges. Importantly, from a process manufacture perspective, if
fiber reinforcement is to be employed, then it is preferable to
intermix and to pre-wet the cellulose fibrous components with a
hydrophilic formulation prior to the onset of appreciable gelation.
This is because the fibers can be more easily wetted and
intermeshed with the formulation when the formulation exhibits low
viscosity and minimal elastic recovery as it would prior to
gelation. In order to accomplish this process step, there needs to
be ample work time prior to gelation to facilitate the total time
requirements for vehicle mixing, metering, wetting, and
infiltration/impregnation of the fibrous material.
[0411] As illustrated in Example 14, the work time window prior to
gelation is significantly shortened for formulations comprising
hydrophilic liquids. For example, when water is mixed with GLBG at
a 2/1 (w/w) ratio, gelation and elastic network formation was
observed to begin almost immediately.
[0412] However, for the case of glycerin, the work time window
prior to the onset of gelation was observed to be significantly
longer, thereby making glycerin a more practical choice as a liquid
for creating more hydrophilic hemostatic fiber-reinforced delivery
devices. It is understood by those skilled in the art that within
some time period after mixing liquids like water or glycerin with
gelatin, gelation will begin to occur, and the initial suspension
of discrete gelatin particulates will become transformed into an
elastic gelled network of surface-bonded, aggregated gelatin
particulates. In the present example, the time-period preceding
gelation, herein referred to as the "work-time") defines the window
of time that enables the product to be made through the process of
impregnating a fibrous substrate. As long as the process is
initiated during the work-time prior to gelation, the viscosity and
elasticity of the vehicle will be low enough to enable facile
impregnation of fibrous substrates with high expediency. Thus, it
is desirable that the gelation process be made to occur after the
fibrous textile is impregnated with the formulation, and not
before.
[0413] For the purposes of creating a more hydrophilic
fiber-reinforced delivery device, it is desirable that the liquid
component be miscible enough with the hygroscopic network-forming
component, including gums like gelatin, gum arabic, ghatti, karaya,
tragacanth, agar, Irish moss, carrageenin, alginates, seed extracts
of which include locust bean, locust kernel, and quince seed gums
as examples, manufactured and modified dextrins and British gums,
water-dispersible or soluble derivatives of cellulose, etc., to
lead to gelation and to the formation of a plasticized polymer
network. It is further desirable that the work-time prior to
gelation be long enough to facilitate all of the process steps that
are required for product formation, such as vehicle mixing,
metering, conveying, wetting, pressing, etc. The work-time window
for textile impregnation can be determined from the suspension test
as defined in Example 14. If a continuous or semi-continuous
process is used to meter and convey the formulation onto a web of
fibrous material, then the web could be optionally conveyed through
a forced air or infrared heated oven to facilitate faster gelation.
Regardless of the use of ovens, once the gelation process is
complete, the resulting impregnated composite can be cut to achieve
the desired geometric size for the application, and then the
resulting delivery device can be packaged for storage prior to
deployment.
[0414] Regarding storage, it is further desirable that the liquid
be biostable, either on its own, or through the incorporation of
preservatives that guard against bacterial growth during periods of
product manufacturing, packaging and storage. It is also desirable
that the liquid lead to formation of a gelled polymer network after
textile impregnation and not before. One example of a liquid that
meets both criteria is glycerin. Other liquids can be used,
including for example, propylene glycol, polyethylene glycols and
polypropylene glycols of various molecular weights, water-based
natural products like honey, polyhydric alcohols and derivatives of
the same, as well as mixtures of any of these types.
[0415] It is also important that the fibrous components of the
composite delivery device be resistant to deterioration, swelling,
or dissolution by a hydrophilic liquid. Surgicel Original (SO)
textiles were determined to be resistant to glycerin. In a separate
experiment, pre-cut SO textiles ((1.8.times.3.8 cm) were separately
drop-coated with glycerin and water. After 24 hours, the
glycerin-coated textile was observed to retain its meshed structure
with no noticeable evidence of dissolution or physical changes,
including no shrinkage or swelling. In a similar test, the SO
textile was also observed to be more resistant to water than its
SafeGauze counterpart. SafeGauze dissolved upon exposure to water
as shown in Example 5, whereas SO showed no apparent signs of
dissolution within a 24-hour window of testing, only shrinkage.
[0416] Regardless of whether a delivery device is designed to be
more hydrophobic or hydrophilic, the resistance of the fibrous
material to water dissolution or to degradation can be an important
and desirable attribute, particularly after deployment of the
delivery device. Although it is desirable that the fibrous material
eventually degrade and become bio-absorbed, it is still desirable
that the fibrous material maintain integrity for a period of time
during the post-deployment lifetime of the device, mainly because
the retention of a composite structure with fibrous reinforcement
is conducive to maximizing macroscopic erosion resistance, which is
another desirable attribute for longer-term durability if the
delivery device is deployed in an oral tooth socket
application.
[0417] In the present example, the following steps were taken to
prepare two composite-reinforced delivery devices with glycerin as
the liquid component in the formulation. Samples 15A and 15B with
compositions are provided in Tables 15-1 and 15-2.
Sample 15A.
[0418] Step-1: a segment of Surgicel Original (SO) oxidized
cellulose textile was cut (1.8.times.3.8 cm) and weighed at 0.0475
g; Step-2: 0.3061 g of PLGA-encapsulated BUP (SWRI; sample
18-0202-015-21; 20% w/w BUP loaded; Resomer RG 504; D50=4.3
microns) was pre-weighed into a 15 ml HDPE beaker; Step-3: a
premixed suspension of Great Lakes bovine gelatin (GLBG) and
glycerin was prepared using 1.8 g GLBG+3.6 g glycerin, and the mix
was allowed to set for 10 minutes; Step-4: 0.4317 g of the premixed
suspension from step 3 was added to the beaker with the pre-weighed
PLGA-encapsulated microspheres, and the resulting vehicle was mixed
by hand for approximately 5 minutes with a spatula until it formed
a homogeneous cream; Step-5: using a spatula, 0.6189 g of the cream
from step-5 was coated and spread over the entire length of a
single pre-weighed textile from step-1, and then the textile was
folded once in its center, over and onto itself before being
subjected to light pressing with the spatula to achieve
impregnation; Step-6: The square shaped impregnated device was
weighed to a final weight of textile+vehicle=0.5852 g, equating to
a final weight after transfer loss=0.5377 g; Step-7: the delivery
device was then allowed to set and gel under ambient conditions (20
degrees C.), and then was qualitatively monitored over time.
[0419] Initially, the more hydrophilic 15A formulation as prepared
in step-4 was noted to be qualitatively similar in viscosity and in
compliance to the comparable, but more hydrophobic formulation of
14C-3A that was prepared in Example 12. After textile impregnation
was completed in step-6, the device was also noted to be
qualitatively similar in stiffness and in compliance to the
analogous delivery device that was prepared in Example 12.
[0420] After approximately 2 hours, the delivery device was still
cohesively intact, but it had become noticeably more stiff owing to
the onset of gelation. Note that the 2/1 glycerin/GLBG (w/w)
mixture that was retained from step-3 had become waxy and higher in
viscosity at this stage. After 15 hours, the 2/1 glycerin/GLBG
(w/w) mixture that was retained from step-3 had become a solid
elastic network. The device itself exhibited internal cohesive
failure as it had opened along its fold to reveal a powdery and
friable surface of cohesively failed formulation. The stress of the
fold in the textile coupled with swelling stresses from the
glycerin-infused gelatin particles was substantial enough to cause
cohesive failure of the gelled mixture.
[0421] Thus, unlike the comparable composite reinforced delivery
device 14C-3A from Example 12, the more hydrophilic delivery device
of sample 15A was unable to retain enough cohesive strength after
gelation to resist swelling stresses and to remain cohesively
intact, thereby illustrating one of the difficulties in
manufacturing a more hydrophilic composite reinforced delivery
device which is gelled prior to deployment. This result serves to
demonstrate one of the limitations of a more hydrophilic delivery
device that does not occur with comparable hydrophobic devices.
Specifically, higher total binder levels (e.g., water+GLBG or
glycerin+GLBG) are required for devices where the binder is
designed to be gelled prior to deployment. This is necessary not
only to provide adequate compliance for deployment, but to also
provide mechanical properties that are commensurate with those
needed to manufacture and store a textile-impregnated delivery
device. Thus, like its water-gelled counterparts as described
earlier in 618-1B from Example 12, this indicates that one of at
least three things would have to be done to create a viable
composition: 1) increase the total binder level (gelatin+glycerin);
2) increase the glycerin/gelatin weight ratio to be more akin to
what was used in more hydrophilic delivery devices with water to
about 4/1 (w/w) instead of 2/1 (w/w); or 3) exercise some
combination of both.
[0422] However, it must be borne in mind that one consequence of
these approaches is that the delivery device and its active
ingredients will become diluted. Of course, the ramifications of
this are dependent on the end use application requirements, and on
the net dosage-level requirements of active ingredients that are
needed for the end use application.
[0423] In accordance with this thinking, sample 15B was prepared
using a 3.92/1 w/w ratio of glycerin to GLBG instead of 2/1 w/w,
and a total vehicle binder level (glycerin+gelatin) of 62% by
weight instead of 58.51% by weight. The steps used in preparing 15B
are provided below.
Sample 15B.
[0424] Step-1: a segment of Surgicel Original (SO) oxidized
cellulose textile was cut (1.8.times.3.8 cm) and weighed at 0.0489
g; Step-2: 0.3064 g of PLGA-encapsulated BUP (SWRI; sample
18-0202-015-21; 20% w/w BUP loaded; Resomer RG 504; D50=4.3
microns) was pre-weighed into a 15 ml HDPE beaker; Step-3: a
premixed suspension of Great Lakes bovine gelatin (GLBG) and
glycerin was prepared using 1.0 g GLBG+3.92 g glycerin, and then
the mix was allowed to set for approximately 20 minutes; Step-4:
0.5 g of the premixed suspension from step 3 was added to the
beaker with the pre-weighed PLGA-encapsulated microspheres, and the
resulting vehicle was mixed by hand with a spatula for
approximately 10 minutes until it formed a homogeneous cream;
Step-5: using a spatula, 0.6119 g of the vehicle cream from step-5
was coated and spread over the entire length of a single
pre-weighed textile from step-1, and then the textile was folded
once in its center, over and onto itself before being subjected to
light pressing with the spatula to achieve impregnation; Step-6:
The square shaped impregnated device was weighed to a final weight
of textile+vehicle=0.5859 g, equating to a final vehicle weight
after transfer loss=0.5370 g; Step-7: the delivery device was then
allowed to set and gel under ambient conditions (20 degrees C.),
and it was qualitatively monitored over time.
[0425] Initially, the 15B formulation as prepared in step-4 was
noted to be qualitatively similar in viscosity to sample 15A at the
same stage of the process. After textile impregnation was completed
in step-6, the delivery device was also noted to be qualitatively
similar in stiffness and in compliance to 15A, and to the analogous
more hydrophobic delivery device that was prepared in Example
12.
[0426] After approximately 2 hours, the delivery device of 15B was
still cohesively intact, but unlike 15A, there was no noticeable
qualitative change in the compliance of the device. Also, unlike
the 15A premix of glycerin and gelatin that had become waxy at this
stage, the 3.92/1 (w/w) premix for 15B was still a pourable
liquid.
[0427] After approximately 15 hours, the 15B delivery device had
become noticeably more stiff owing to the onset of gelation, but it
was cohesively intact. At this stage, its stiffness was
qualitatively similar to that of 15A at time=2 hours. Similarly,
the 3.92/1 (w/w) premix from 15B was waxy, much like the 15A premix
had appeared after only two hours. By contrast, the 2/1 (w/w) 15A
premix had become an elastic network at t=15 hours.
[0428] After approximately 24 hours, the 15B device did not exhibit
a noticeable change, and it was still cohesively intact. In
addition, the 3.92/1 (w/w) premix from 15B had become noticeably
more elastic.
[0429] The 15B delivery device continued to remain mechanically
stable and unchanged throughout the duration of the experiment of
48 hours.
[0430] The compositions of the 15A and 15B vehicles are provided in
Table 15-1, and the final device compositions are provided in Table
15-2. Note that the level of dispersed solids is expressed for two
different physical states of the formulations--before gelation,
while glycerin is the continuous phase for dispersed particulates
of gelatin and PLGA), and after gelation when plasticized gelatin
becomes the continuous phase for the dispersion of PLGA.
TABLE-US-00037 TABLE 15-1 Weight % compositions of hydrophilic
vehicles for use in preparing textile-impregnated devices made with
glycerin as the liquid carrier for the vehicle. Calculations also
include the net weight % concentration of BUP in each vehicle, the
net PLGA polymer weight % (i.e., ~80% of the weight of BUP loaded
microspheres), the total weight % of dispersed solids in the
vehicle prior to gelation, and the total weight % of dispersed
solids in the matrix after gelation (the continuous phase is
glycerin prior to gelation, and plasticized glycerin after
gelation). Vehicle Mixture Composition 15A 15B glycerin 39.01%
49.4% Bovine Gelatin 19.50% 12.6% 5 um PLGA Placebo microspheres 0%
0% 4.3 micron 20% BUP free base loaded 41.49% 38.0% PLGA
microspheres BUP free base (directly added to vehicle) 0% 0% TOTAL
100.00% 100% Total BUP in vehicle 8.30% 7.6% Total PLGA polymer in
Vehicle 33.19% 30.4% Total % dispersed solids in liquid vehicle
60.99% 50.6% prior to gelation = 100 .times. (PLGA-BUP +
GLBG)/(glycerin + GLBG + PLGA-BUP) Total % dispersed solids in
gelled vehicle 41.49% .sup. 38% matrix phase = 100 .times. (PLGA-
BUP)/(glycerin + GLBG + PLGA-BUP)
TABLE-US-00038 TABLE 15-2 Weight % compositions of hydrophilic
textile-impregnated devices made with glycerin as the liquid
carrier for the vehicle. The vehicle compositions as reported in
Table 15-1 were separately impregnated into individual SO textiles.
The calculations for compositions also include the weight %
concentration of BUP, and the effective available BUP concentration
for release on a unit weight of device basis (mg/g). Ingredient 15A
15B Great Lakes Bovine Gelatin (GLBG) 17.92% 11.55% glycerin 35.84%
45.28% PLGA polymer (i.e., representing 30.50% 27.86% 80% of the
weight of 4.3-micron microspheres loaded with 20% by wt. BUP)
Encapsulated BUP (i.e., representing 7.62% 6.96% 20% by weight of
the 4.3-micron microspheres loaded with 20% by wt. BUP) PLGA
polymer from 5-micron 0% 0% placebo microspheres BUP free base
(non-encapsulated, 0% 0% directly added to the vehicle) SO textile
8.12% 8.35% mg BUP/g device 76 70 Weight of Device as made (g)
0.5852 0.5859 Weight of Vehicle as made (g) 0.5377 0.5370
Example 16. Preparation of a Temperature Activated Hydrophobic
Device with Coconut Oil as the Liquid Component
[0431] A delivery device analogous to 14C-3A from Example 12 was
prepared using coconut oil (CO) as the liquid carrier in place of
mineral oil. The CO as discussed in Example 14 was deemed to be
suitable for use as a liquid carrier in preparing a more
hydrophobic device. The compositions of the vehicle and the device
are provided in Tables 16-1 and 16-2. The formulation and textile
impregnated delivery devices were prepared using procedures
outlined in Examples 9, 12 and 13. However, while preparing both
the premix of gelatin with CO and the formulation with added PLGA,
the temperature was maintained at 27 degrees C., which is above the
melt point of the CO. The CO/gelatin premix was observed to
solidify upon cooling to 20 degrees C. This process of
solidification and melting was observed to be reversible for both
the premix, and for the resulting formulation. While in its liquid
dispersion state, the formulation was coated onto a precut SO
textile. Initially, at 27 degrees C., it was a compliant device,
qualitatively similar in compliance characteristics to the
analogous device prepared with mineral oil in Example 12 (14C-3A).
When the device was allowed to cool to 20 degrees C., it became
noticeably stiffer. Upon re-heating to 27 degrees C., it became
noticeably compliant again like its 14C-3A counterpart. This
process was observed to be reversible over multiple cycles.
[0432] CO is a complex mixture of symmetric and asymmetric
triglycerides. Some of the components within the CO have melt
points that render the mixture as having the capability of
exhibiting solid-like characteristics at 20 degrees C. and liquid
characteristics at 27 degrees C. Importantly, it is possible to
formulate any oil that is deemed to be suitable for use in a more
hydrophobic device with waxes, fatty acid esters, or mixtures
thereof at appropriate weight ratios to create carriers with melt
points that can be tuned to any temperature, including body
temperature. In so doing, a temperature activated device can be
made to soften and or to harden at specific temperatures, thereby
changing its mechanical characteristics or time-release
characteristics.
[0433] By using these teachings, a delivery device can be made to
soften at or above 37 degrees C. (body temperature) and to freeze
or harden when it becomes cooled. The advantage is that through
pre-heating the device, it can be made conformable for optimal
placement into the tooth extraction socket. Upon cooling to body
temperature, the delivery device can then be made to harden via
recrystallization of components that have been formulated into the
vehicle.
[0434] Conversely, a delivery device can be made to soften upon
deployment. This can be accomplished by tuning the melt point of
the vehicle to be near or below body temperature.
[0435] Either of these approaches can have an impact on end use
characteristics. For example, a softer and more compliant delivery
device is easier to conform to the geometric shape of a cavity. A
delivery device with higher modulus can exhibit better resistance
to erosion. For example, a delivery device that remains soft and
conformable after deployment could be made to temporarily harden if
the patient consumes a cold liquid. This can result in improved
erosion resistance on-demand upon exposure to the cooler liquid as
it flows across an exposed surface of the delivery device.
[0436] Release rates and fluid influx rates will also be affected
by the compliance of the delivery device, with diffusion being
slower through a more rigid matrix medium than through a softer
medium.
TABLE-US-00039 TABLE 16-1 Weight % compositions of a hydrophobic
vehicle for use in preparing textile-impregnated devices using
coconut oil as the liquid carrier for the vehicle. Calculations
also include the net weight % concentration of BUP in each vehicle,
the net PLGA polymer weight % (i.e., ~80% of the weight of BUP
loaded microspheres), and the total weight % of dispersed solids.
Vehicle Mixture Composition 16A Coconut Oil 39.01% Beeswax 0%
Bovine Gelatin 19.50% 5 um PLGA Placebo microspheres 0% 4.3 micron
20% BUP free base loaded PLGA microspheres 41.49% BUP free base
(directly added to vehicle) 0% TOTAL 100.00% Total BUP in vehicle
8.30% Total PLGA polymer in Vehicle 33.19% Total % dispersed solids
in vehicle 60.99%
TABLE-US-00040 TABLE 16-2 Weight % composition of a hydrophobic
textile-impregnated device made with coconut oil. The vehicle
composition as reported in Table 16-1 was impregnated into a single
SO textile to yield the empirical composition as presented below.
The calculations also include the weight % concentration of BUP,
and the effective available BUP concentration for release per unit
weight of device (mg/g). Ingredient 16A Great Lakes Bovine Gelatin
(GLBG) 15.32% Coconut Oil (CO) 30.63% Beeswax (BW) 0% PLGA polymer
(i.e., representing 80% of the weight of 36.76% 4.3-micron
microspheres loaded with 20% by wt. BUP) Encapsulated BUP (i.e.,
representing 20% by weight of 9.19% the 4.3-micron microspheres
loaded with 20% by wt. BUP) PLGA polymer from 5-micron placebo
microspheres 0% BUP free base (non-encapsulated, directly added to
the 0% vehicle) SO textile 8.10% mg BUP/g device 92 Weight of
Device as made (g) 0.6048 Weight of Vehicle as made (g) 0.5558
* * * * *